Devices for reducing left atrial pressure, and methods of making and using same

ABSTRACT

A device for regulating blood pressure between a patient&#39;s left atrium and right atrium comprises an hourglass-shaped stent comprising a neck region and first and second flared end regions, the neck region disposed between the first and second end regions and configured to engage the fossa ovalis of the patient&#39;s atrial septum; and a one-way tissue valve coupled to the first flared end region and configured to shunt blood from the left atrium to the right atrium when blood pressure in the left atrium exceeds blood pressure in the right atrium. The inventive devices may reduce left atrial pressure and left ventricular end diastolic pressure, and may increase cardiac output, increase ejection fraction, relieve pulmonary congestion, and lower pulmonary artery pressure, among other benefits. The inventive devices may be used, for example, to treat subjects having heart failure, pulmonary congestion, or myocardial infarction, among other pathologies.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.14/712,801, filed May 14, 2015, now U.S. Pat. No. 9,980,815, which is adivisional of U.S. patent application Ser. No. 13/193,335, filed Jul.28, 2011, now U.S. Pat. No. 9,034,034, which claims the benefit of U.S.Provisional Patent Application Ser. No. 61/425,792, filed Dec. 22, 2010,the entire contents of which are incorporated by reference herein, andU.S. patent application Ser. No. 13/193,335 is also acontinuation-in-part of International Patent Application No.PCT/IL2010/000354, filed May 4, 2010, which claims the benefit of U.S.Provisional Patent Application Nos. 61/175,073, filed May 4, 2009 and61/240,667, filed Sep. 9, 2009, the entire contents of each of which areincorporated by reference herein.

FIELD OF THE INVENTION

This application generally relates to devices and methods for reducingleft atrial pressure, particularly in subjects with heart pathologiessuch as congestive heart failure (CHF) or myocardial infarction (MI).

BACKGROUND OF THE INVENTION

Heart failure is the physiological state in which cardiac output isinsufficient to meet the needs of the body and the lungs. CHF occurswhen cardiac output is relatively low and the body becomes congestedwith fluid. There are many possible underlying causes of CHF, includingmyocardial infarction, coronary heart disease, valvular disease, andmyocarditis. Chronic heart failure is associated with neurohormonalactivation and alterations in autonomic control. Although thesecompensatory neurohormonal mechanisms provide valuable support for theheart under normal physiological circumstances, they also have afundamental role in the development and subsequent progression of CHF.For example, one of the body's main compensatory mechanisms for reducedblood flow in CHF is to increase the amount of salt and water retainedby the kidneys. Retaining salt and water, instead of excreting it intothe urine, increases the volume of blood in the bloodstream and helps tomaintain blood pressure. However, the larger volume of blood alsostretches the heart muscle, enlarging the heart chambers, particularlythe ventricles. At a certain amount of stretching, the heart'scontractions become weakened, and the heart failure worsens. Anothercompensatory mechanism is vasoconstriction of the arterial system. Thismechanism, like salt and water retention, raises the blood pressure tohelp maintain adequate perfusion.

In low ejection fraction (EF) heart failure, high pressures in the heartresult from the body's attempt to maintain the high pressures needed foradequate peripheral perfusion. However, the heart weakens as a result ofthe high pressures, aggravating the disorder. Pressure in the leftatrium may exceed 25 mmHg, at which stage, fluids from the blood flowingthrough the pulmonary circulatory system flow out of the interstitialspaces and into the alveoli, causing pulmonary edema and lungcongestion.

Table 1 lists typical ranges of right atrial pressure (RAP), rightventricular pressure (RVP), left atrial pressure (LAP), left ventricularpressure (LVP), cardiac output (CO), and stroke volume (SV) for a normalheart and for a heart suffering from CHF. In a normal heart beating ataround 70 beats/minute, the stroke volume needed to maintain normalcardiac output is about 60 to 100 milliliters. When the preload,after-load, and contractility of the heart are normal, the pressuresrequired to achieve normal cardiac output are listed in Table 1. In aheart suffering from CHF, the hemodynamic parameters change (as shown inTable 1) to maximize peripheral perfusion.

TABLE 1 Parameter Normal Range CHF Range RAP (mmHg) 2-6  6-15 RVP (mmHg)15-25 20-40 LAP (mmHg)  6-12 15-30 LVP (mmHg)  6-120  20-220 CO(liters/minute) 4-8 2-6 SV (milliliters/beat)  60-100 30-80

CHF is generally classified as either systolic heart failure (SHF) ordiastolic heart failure (DHF). In SHF, the pumping action of the heartis reduced or weakened. A common clinical measurement is the ejectionfraction, which is a function of the blood ejected out of the leftventricle (stroke volume), divided by the maximum volume remaining inthe left ventricle at the end of diastole or relaxation phase. A normalejection fraction is greater than 50%. Systolic heart failure has adecreased ejection fraction of less than 50%. A patient with SHF mayusually have a larger left ventricle because of a phenomenon calledcardiac remodeling that occurs secondarily to the higher ventricularpressures.

In DHF, the heart generally contracts normally, with a normal ejectionfraction, but is stiffer, or less compliant, than a healthy heart wouldbe when relaxing and filling with blood. This stiffness may impede bloodfrom filling the heart, and produce backup into the lungs, which mayresult in pulmonary venous hypertension and lung edema. DHF is morecommon in patients older than 75 years, especially in women with highblood pressure.

Both variants of CHF have been treated using pharmacological approaches,which typically involve the use of vasodilators for reducing theworkload of the heart by reducing systemic vascular resistance, as wellas diuretics, which inhibit fluid accumulation and edema formation, andreduce cardiac filling pressure.

In more severe cases of CHF, assist devices such as mechanical pumpshave been used to reduce the load on the heart by performing all or partof the pumping function normally done by the heart. Chronic leftventricular assist devices (LVAD), and cardiac transplantation, oftenare used as measures of last resort. However, such assist devices aretypically intended to improve the pumping capacity of the heart, toincrease cardiac output to levels compatible with normal life, and tosustain the patient until a donor heart for transplantation becomesavailable. Such mechanical devices enable propulsion of significantvolumes of blood (liters/min), but are limited by a need for a powersupply, relatively large pumps, and the risk of hemolysis, thrombusformation, and infection. Temporary assist devices, intra-aorticballoons, and pacing devices have also been used.

In addition to cardiac transplant, which is highly invasive and limitedby the ability of donor hearts, surgical approaches such as dynamiccardiomyoplastic or the Batista partial left ventriculectomy may also beused in severe cases.

Various devices have been developed using stents or conduits to modifyblood pressure and flow within a given vessel, or between chambers ofthe heart. For example, U.S. Pat. No. 6,120,534 to Ruiz is directed toan endoluminal stent for regulating the flow of fluids through a bodyvessel or organ, for example for regulating blood flow through thepulmonary artery to treat congenital heart defects. The stent mayinclude an expandable mesh having lobed or conical portions joined by aconstricted region, which limits flow through the stent. The mesh maycomprise longitudinal struts connected by transverse sinusoidal orserpentine connecting members. Ruiz is silent on the treatment of CHF orthe reduction of left atrial pressure.

U.S. Pat. No. 6,468,303 to Amplatz et al. discloses a collapsiblemedical device and associated method for shunting selected organs andvessels. Amplatz discloses that the device may be suitable to shunt aseptal defect of a patient's heart, for example, by creating a shunt inthe atrial septum of a neonate with hypoplastic left heart syndrome(HLHS). Amplatz discloses that increasing mixing of pulmonary andsystemic venous blood improves oxygen saturation. Amplatz discloses thatdepending on the hemodynamics, the shunting passage can later be closedby an occluding device. Amplatz is silent on the treatment of CHF or thereduction of left atrial pressure, as well as on means for regulatingthe rate of blood flow through the device.

U.S. Patent Publication No. 2005/0165344 to Dobak, III discloses anapparatus for treating heart failure that includes a conduit positionedin a hole in the atrial septum of the heart, to allow flow from the leftatrium into the right atrium. Dobak discloses that the shunting of bloodwill reduce left atrial pressures, thereby preventing pulmonary edemaand progressive left ventricular dysfunction, and reducing LVEDP. Dobakdiscloses that the conduit may include a self-expandable tube withretention struts, such as metallic arms that exert a slight force on theatrial septum on both sides and pinch or clamp the valve to the septum,and a one-way valve member, such as a tilting disk, bileaflet design, ora flap valve formed of fixed animal pericardial tissue. However, Dobakstates that a valved design may not be optimal due to a risk of bloodstasis and thrombus formation on the valve, and that valves can alsodamage blood components due to turbulent flow effects. Dobak does notprovide any specific guidance on how to avoid such problems.

SUMMARY OF THE INVENTION

Embodiments of the present invention provide hourglass-shaped devicesfor reducing left atrial pressure, and methods of making and using thesame. As elaborated further herein, such reductions in left atrialpressure may increase cardiac output, relieve pulmonary congestion, andlower pulmonary artery pressure, among other benefits. The inventivedevices are configured for implantation through the atrial septum, andparticularly through the middle of the fossa ovalis, away from thesurrounding limbus, inferior vena cava (IVC), and atrial wall. Thedevices are configured to provide one-way blood flow from the leftatrium to the right atrium when the pressure in the left atrium exceedsthe pressure in the right atrium, and thus decompress the left atrium.Such a lowering of left atrial pressure may offset abnormal hemodynamicsassociated with CHF, for example, to reduce congestion as well as theoccurrence of acute cardiogenic pulmonary edema (ACPE), which is asevere manifestation of CHF in which fluid leaks from pulmonarycapillaries into the interstitium and alveoli of the lung. Inparticular, lowering the left atrial pressure may improve the cardiacfunction by:

(1) Decreasing the overall pulmonary circulation pressure, thusdecreasing the afterload on the heart,

(2) Increasing cardiac output by reducing left ventricular end systolicdimensions, and

(3) Reducing the left ventricular end-diastolic pressure (LVEDP) andpulmonary artery pressure (PAP), which in turn may enable the heart towork more efficiently and over time increase cardiac output. Forexample, the oxygen uptake of the myocardium may be reduced, creating amore efficient working point for the myocardium.

As described in further detail below, the devices provided hereincomprise an hourglass or “diabolo” shaped stent encapsulated with abiocompatible material, and secured (e.g., sutured) to a tissue valve.The stent, which may be formed of shape memory material, for example ashape memory metal such as NiTi, comprises a neck region disposedbetween two flared end regions. The tissue valve is coupled to a flaredend region configured for implantation in the right atrium.Specifically, the device may be implanted by forming a puncture throughthe atrial septum, particularly through the fossa ovalis, and thenpercutaneously inserting the device therethrough such that the necklodges in the puncture, the flared end to which the tissue valve iscoupled engages the right side of the atrial septum, and the otherflared end flanks the left side of the atrial septum (e.g., is spacedapart from and does not contact the left side of the atrial septum).Placement in the middle of the fossa ovalis is useful because theengagement of the right-side flared end with the atrial septum enhancesthe stability of the valve. The neck region and the flared end regionfor placement in the left atrium may each be covered with abiocompatible polymer, such as expanded polytetrafluoroethylene (ePTFE),polyurethane, DACRON (polyethylene terephthalate), silicone,polycarbonate urethane, or pericardial tissue from an equine, bovine, orporcine source, which is optionally treated so as to promote a limitedamount of tissue ingrowth, e.g., of epithelial tissue or a neointimalayer. The tissue valve is connected to the biocompatible polymer in theright-side flared end region, close to the neck region, and ispreferably a tricuspid, bicuspid, or duckbill valve configured to allowblood to flow from the left atrium to the right atrium when the pressurein the left atrium exceeds that in the right atrium, but prevent flowfrom the right atrium to the left atrium. In preferred embodiments, thedevice is effective to maintain the pressure differential between theleft atrium and right atrium to 15 mmHg or less.

Under one aspect of the present invention, a device for regulating bloodpressure between a patient's left atrium and right atrium comprises anhourglass-shaped stent comprising a neck and first and second flared endregions, the neck disposed between the first and second end regions andconfigured to engage the fossa ovalis of the patient's atrial septum;and a one-way tissue valve coupled to the first flared end region andconfigured to shunt blood from the left atrium to the right atrium whenblood pressure in the left atrium exceeds blood pressure in the rightatrium. In accordance with one aspect of the invention, moving portionsof the valve are disposed in the right atrium, joined to but spacedapart from the neck region.

The hourglass-shaped stent may include a shape memory material (e.g.,metal) coated with a biocompatible polymer from a portion of the firstflared end region, through the neck region, and through the secondflared end region, and the tissue valve may extend between the firstflared end region and the biocompatible polymer. Providing the tissuevalve in the side of the device to be implanted in the right atrium(that is, in the first flared end region) may inhibit thrombus formationand tissue ingrowth by providing that the tissue valve, as well as theregion where the tissue valve is secured (e.g., sutured) to thebiocompatible polymer, is continuously flushed with blood flowingthrough the right atrium. By comparison, if the tissue valve was insteadsecured (e.g., sutured) to the biocompatible polymer in the neck region,then the interface between the two would contact the tissue of the fossaovalis, which potentially would encourage excessive tissue ingrowth,create leakages, and cause inflammation. Moreover, tissue ingrowth intothe neck region would cause a step in the flow of blood in the narrowestpart of the device, where flow is fastest, which would increase shearstresses and cause coagulation. Instead providing the tissue valveentirely within the right atrial side of the device inhibits contactbetween the tissue valve and the tissue of the atrial septum and fossaovalis. Further, any tissue that ingrows into the valve will notsubstantially affect blood flow through the device, because the valve islocated in a portion of the device having a significantly largerdiameter than the neck region. Moreover, if the biocompatible tissuewere instead to continue on the portions of the frame positioned overthe tissue valve, it may create locations of blood stasis between theleaflets of the tissue valve and the biocompatible material. Having thevalve entirely on the right atrial side and without biocompatiblematerial on the overlying frame enables continuous flushing of theexternal sides of the tissue valve with blood circulating in the rightatrium.

The biocompatible material preferably promotes limited (or inhibitsexcessive) tissue ingrowth into the valve, the tissue ingrowth includingan endothelial layer or neointima layer inhibiting thrombogenicity ofthe device. The endothelial or neointima layer may grow to a thicknessof 0.2 mm or less, so as to render the material inert and inhibithyperplasia.

The hourglass-shaped stent may include a plurality of sinusoidal ringsinterconnected by longitudinally extending struts. In some embodiments,when the shunt is deployed across the patient's atrial septum, the firstflared end region protrudes 5.5 to 7.5 mm into the right atrium. Thesecond flared end region may protrude 2.5 to 7 mm into the left atrium.The neck may have a diameter of 4.5 to 5.5 mm. The first flared endregion may have a diameter between 9 and 13 mm, and the second flaredend region may have a diameter between 8 and 15 mm. The first and secondflared end regions each may flare by about 50 to 120 degrees. Forexample, in one embodiment, the first flared end region flares by about80 degrees, that is, the steepest part of the outer surface of the firstflared end region is at an angle of approximately 40 degrees relative toa central longitudinal axis of the device. The second flared end regionmay flare by about 75 degrees, that is, the steepest part of the outersurface of the second flared end region may be at an angle ofapproximately 37.5 degrees relative to the central longitudinal axis ofthe device.

The inlet of the tissue valve may be about 1-3 mm from a narrowestportion of the neck region, and the outlet of the tissue valve may beabout 5-8 mm from the narrowest portion of the neck region. The tissuevalve may comprise a sheet of tissue having a flattened length of about10-16 mm, and the sheet of tissue may be folded and sutured so as todefine two or more leaflets each having a length of about 5-8 mm. Forexample, the tissue sheet may have a flattened length of no greater than18 mm, for example, a length of 10-16 mm, or 12-14 mm, or 14-18 mm, andmay be folded and sutured to define two or more leaflets each having alength of, for example, 9 mm or less, or 8 mm or less, or 7 mm or less,or 6 mm or less, or even 5 mm or less, e.g., 5-8 mm. The tissue sheetmay have a flattened height no greater than 10 mm, for example, a heightof 2-10 mm, or 4-10 mm, or 4-8 mm, or 6-8 mm, or 4-6 mm. The tissuesheet may have a flattened area of no greater than 150 square mm, forexample, 60-150 square mm, or 80-120 square mm, or 100-140 square mm, or60-100 square mm.

The hourglass-shaped stent may be configured to transition between acollapsed state suitable for percutaneous delivery and an expanded statewhen deployed across the patient's fossa ovalis. The stent may have anhourglass configuration in the expanded state. The hourglassconfiguration may be asymmetric. The stent may be configured forimplantation through the middle of the fossa ovalis, away from thesurrounding limbus, inferior vena cava, and atrial wall.

The one-way tissue valve may have two or more leaflets, e.g., may have atricuspid or bicuspid design. The one-way tissue valve may comprisepericardial tissue, which in one embodiment may consist primarily of themesothelial and loose connective tissue layers, and substantially nodense fibrous layer. Note that the dimensions of the hourglass-shapeddevice may be significantly smaller than those of replacement aorticvalves, which may for example have a diameter of 23 mm and require theuse of larger, thicker valve leaflets to maintain the higher stressesgenerated by the combination of higher pressures and larger diameters.By comparison, the inventive device has much smaller dimensions,allowing the use of thinner tissue (e.g., about one third the thicknessof tissue used in a replacement aortic valve), for example, pericardialtissue in which the external dense fibrous layer is delaminated and themesothelial and loose connective tissue is retained.

Under another aspect of the present invention, a device for regulatingblood pressure between a patient's left atrium and right atrium includesa stent comprising a neck region and first and second flared endregions, the neck region disposed between the first and second endregions and configured to engage the fossa ovalis of the patient'satrial septum; a biocompatible material disposed on the stent in theneck and the second flared end region and a portion of the first flaredend region; and a one-way tissue valve configured to shunt blood fromthe left atrium to the right atrium when blood pressure in the leftatrium exceeds blood pressure in the right atrium, the valve having anoutlet coupled to the first flared end region and an inlet coupled to anedge of the biocompatible material, the valve and the biocompatiblematerial defining a continuous sheath that inhibits excessive tissueingrowth into the valve and channels blood flow through the valve. Inone embodiment, the edge of the biocompatible material is about 1-3 mm,e.g., 2 mm, from a narrowest portion of the neck region.

Under another aspect, a method of treating a subject with heartpathology comprises: providing a device having first and second flaredend regions and a neck region disposed therebetween, and a tissue valvecoupled to the first flared end region; deploying the device across apuncture through the subject's fossa ovalis such that the neck region ispositioned in the puncture, the first flared end region is disposed in,and engages, the atrial septum, and the second flared end region isdisposed in, and flanks, the atrial septum; and reducing left atrialpressure and left ventricular end diastolic pressure by shunting bloodfrom the left atrium to the right atrium through the device when theleft atrial pressure exceeds the right atrial pressure.

Subjects with a variety of heart pathologies may be treated with, andmay benefit from, the inventive device. For example, subjects withmyocardial infarction may be treated, for example by deploying thedevice during a period immediately following the myocardial infarction,e.g., within six months after the myocardial infarction, or within twoweeks following the myocardial infarction. Other heart pathologies thatmay be treated include heart failure and pulmonary congestion. Reducingthe left atrial pressure and left ventricular end diastolic pressure mayprovide a variety of benefits, including but not limited to increasingcardiac output; decreasing pulmonary congestion; decreasing pulmonaryartery pressure; increasing ejection fraction; increasing fractionalshortening; and decreasing left ventricle internal diameter in systole.Means may be provided for measuring such parameters.

Such methods may include identifying the middle of the fossa ovalis ofthe atrial septum by pressing a needle against the fossa ovalis topartially tent the fossa ovalis; and puncturing the middle of the fossaovalis with the needle.

Under yet another aspect of the present invention, a method of making adevice comprises: providing a tube of shape-memory metal; expanding thetube on a mandrel to define first and second flared end regions and aneck therebetween, and heating the expanded tube to set the shape;coating the neck and second flared end region with a biocompatiblematerial; providing a valve of animal pericardial tissue having leafletsfixed in a normally closed position; and securing an inlet of the valveto the first flared end region and to the biocompatible polymer at theneck region. The tube may be laser cut and may include a plurality ofsinusoidal rings connected by longitudinally extending struts, and thevalve may be sutured to the struts and to the biocompatible material toform a passage for blood.

BRIEF DESCRIPTION OF DRAWINGS

FIGS. 1A-1D illustrate perspective views of an hourglass-shaped devicehaving a tricuspid valve, according to some embodiments of the presentinvention.

FIG. 2A schematically illustrates a plan view of the right atrial sideof the atrial septum, including a site for implanting anhourglass-shaped device through the middle of the fossa ovalis.

FIG. 2B schematically illustrates a cross-sectional view of thehourglass-shaped device of FIGS. 1A-1D positioned in the fossa ovalis ofthe atrial septum, according to some embodiments of the presentinvention.

FIG. 3A is a flow chart of steps in a method of making anhourglass-shaped device, according to some embodiments of the presentinvention.

FIGS. 3B-3E illustrate plan views of sheets of material for use inpreparing tissue valves, according to some embodiments of the presentinvention.

FIG. 4 is a flow chart of steps in a method of percutaneously implantingan hourglass-shaped device in a puncture through the fossa ovalis,according to some embodiments of the present invention.

FIGS. 5A-5D schematically illustrate steps taken during the method ofFIG. 4, according to some embodiments of the present invention.

FIG. 6A is an image from a computational fluid dynamic model of flowthrough an hourglass-shaped device in the open configuration.

FIG. 6B is a plot showing the relationship between the left-to-rightatrial pressure differential and the flow rate through the valve forhourglass-shaped devices having different valve diameters, according tosome embodiments of the present invention.

FIG. 7 is a flow chart of steps in a method of noninvasively determiningleft atrial pressure using an hourglass-shaped device, and adjusting atreatment plan based on same, according to some embodiments of thepresent invention.

FIGS. 8A-8C illustrate perspective views of an alternativehourglass-shaped device, according to some embodiments of the presentinvention.

FIG. 9 is a perspective view of a further alternative hourglass-shapeddevice, according to some embodiments of the present invention.

FIGS. 10A-10D are plots respectively showing the left atrial pressure,right atrial pressure, ejection fraction, and pulmonary artery pressurein animals into which an exemplary hourglass-shaped device wasimplanted, as well as control animals, during a twelve-week study.

FIGS. 11A-11B are photographic images showing an hourglass-shaped devicefollowing explanation from an animal after being implanted for 12 weeks.

FIG. 11C is a microscope image of a cross-section of an hourglass-shapeddevice following explanation from an animal after being implanted for 12weeks.

DETAILED DESCRIPTION

Embodiments of the present invention are directed to devices that reduceleft atrial pressure, and thus may be useful in treating subjectssuffering from congestive heart failure (CHF) or other disordersassociated with elevated left atrial pressure. Specifically, theinventive device includes an hourglass or “diabolo” shaped stent,preferably formed of a shape memory metal, and a biocompatible valvecoupled thereto. The stent is configured to lodge securely in the atrialseptum, preferably the fossa ovalis, and the valve is configured toallow one-blood flow from the left atrium to the right atrium,preferably through the fossa ovalis, when blood pressure in the leftatrium exceeds that on the right. Usefully, the inventive devices areconfigured so as to reduce blood pressure in the left atrium even whenthe pressure differential therebetween is relatively low; to provide asmooth flow path with a large valve opening, thus inhibiting turbulenceand high shear stresses that would otherwise promote thrombus formation;to seal securely with rapid valve closure when the left and right atrialpressures equalize or the right atrial pressure exceeds left atrialpressure; and to have a relatively small implantation footprint so as toinhibit tissue overgrowth and inflammatory response.

First, a preferred embodiment of the inventive hourglass-shaped devicewill be described, and then methods of making, implanting, and using thesame will be described. Then, the hemodynamic flow characteristics ofsome illustrative devices will be described, as well as a method forusing an hourglass-shaped device to noninvasively determine left atrialpressure based on images of blood flowing through the implanted device.Some alternative embodiments will then be described. Lastly, an Examplewill be provided that describes a study performed on several animalsinto which an exemplary device was implanted, as compared to a group ofcontrol animals.

FIGS. 1A-1D illustrate perspective views of an illustrative embodimentof the inventive device. First, with reference to FIG. 1A, device 100includes an hourglass-shaped stent 110 and tissue valve 130,illustratively, a tricuspid valve including three coapting leaflets.Device 100 has three general regions: first flared or funnel-shaped endregion 102, second flared or funnel-shaped end region 106, and neckregion 104 disposed between the first and second flared end regions.Neck region 104 is configured to lodge in a puncture formed in theatrial septum, preferably in the fossa ovalis, using methods describedin greater detail below. First flared end region 102 is configured toengage the right side of the atrial septum, and second flared end region106 is configured to flank the left side of the atrial septum, whenimplanted. The particular dimensions and configurations of neck region104 and first and second flared end regions 102, 106 may be selected soas to inhibit the formation of eddy currents when implanted, and thusinhibit thrombus formation; to inhibit tissue ingrowth in selectedregions; to promote tissue ingrowth in other selected regions; and toprovide a desirable rate of blood flow between the left and right atria.

Hourglass-shaped stent 110 is preferably formed of a shape memory metal,e.g., NITINOL, or any other suitable material known in the art. Stent110 includes a plurality of sinusoidal rings 112-116 interconnected bylongitudinally extending struts 111. Rings 112-116 and struts 111 may beof unitary construction, that is, entire stent 110 may be laser cut froma tube of shape memory metal. As can be seen in FIG. 1A, neck region 104and second flared end region 106 are covered with biocompatible material120, for example a sheet of a polymer such as expandedpolytetrafluoroethylene (ePTFE), silicone, polycarbonate urethane,DACRON (polyethylene terephthalate), or polyurethane, or of a naturalmaterial such as pericardial tissue, e.g., from an equine, bovine, orporcine source. Specifically, the region extending approximately fromsinusoidal ring 113 to sinusoidal ring 116 is covered with biocompatiblematerial 120. Material 120 preferably is generally smooth so as toinhibit thrombus formation, and optionally may be impregnated withcarbon so as to promote tissue ingrowth. Preferably, portions of stent110 associated with first flared end region 102 are not covered with thebiocompatible material, but are left as bare metal, so as to inhibit theformation of stagnant flow regions in the right atrium that otherwiseand to provide substantially free blood flow around leaflets 131, so asto inhibit significant tissue growth on leaflets 131. The bare metalregions of stent 110, as well as any other regions of the stent,optionally may be electropolished or otherwise treated so as to inhibitthrombus formation, using any suitable method known in the art.

An inlet end of tissue valve 130 is coupled to stent 110 in first flaredend region 102. In the illustrated embodiment, tissue valve 130 is atricuspid valve that includes first, second, and third leaflets 131defining valve opening 132. Other embodiments, illustrated furtherbelow, may include a bicuspid or duckbill valve, or other suitable valveconstruction. However, it is believed that tricuspid valves may provideenhanced leaflet coaptation as compared to other valve types, such thateven if the tissue valve stiffens as a result of tissue ingrowthfollowing implantation, there may still be sufficient leaflet materialto provide coaptation with the other leaflets and close the valve.Preferably, tissue valve 130 opens at a pressure of less than 1 mm Hg,closes at a pressure gradient of between 0-0.5 mm Hg, and remains closedat relatively high back pressures, for example at back pressures of atleast 40 mm Hg. Tissue valve 130 may be formed using any natural orsynthetic biocompatible material, including but not limited topericardial tissue, e.g., bovine, equine, or porcine tissue, or asuitable polymer. Pericardial tissue, and in particular bovinepericardial tissue, is preferred because of its strength and durability.The pericardial tissue may be thinned to enhance compliance, for exampleas described in greater detail below, and may be fixed using anysuitable method, for example, using glutaraldehyde or otherbiocompatible fixative.

As shown in FIG. 1B, tissue valve 130 is coupled, e.g., sutured, tofirst, second, and third longitudinally extending struts 111′, 111″, and111′″ in the region extending between first (uppermost) sinusoidal ring112 and second sinusoidal ring 113. Referring to FIGS. 1A and 1D, tissuevalve 130 is also coupled to the upper edge of biocompatible material120, at or near sinusoidal ring 113, for example along line 121 asshown. As such, tissue valve 130 and biocompatible material 120 togetherprovide a smooth profile to guide blood flow from the left atrium to theright atrium, that is, from the second flared end region 106, throughneck region 104, and through first flared end region 102. In accordancewith one aspect of the invention, the inlet to tissue valve 130 isanchored to neck region 104, such that leaflets 131 extend into theright atrium. In this manner, any eccentricities that may arise from theout-of-roundness of the puncture through the fossa ovalis duringimplantation will not be transferred to the free ends of leaflets 131,thus reducing the risk that any eccentricity of the stent in neck region104 could disturb proper coaptation of the valve leaflets.

FIGS. 1A and 1B illustrate device 100 when tissue valve 130 is in anopen configuration, in which leaflets 131 are in an open position topermit flow, and FIG. 1C illustrates device 100 when tissue valve 130 isin a closed configuration, in which leaflets 131 are in a closedposition to inhibit flow. Tissue valve 130 is configured to open whenthe pressure at second flared end region 106 exceeds that at firstflared end region 102. Preferably, however, tissue valve 130 isconfigured to close and therefore inhibit flow in the oppositedirection, i.e., to inhibit flow from first flared end region 102,through neck region 104, and through second flared end region 104, whenthe pressure at the first flared end region exceeds that of the second.Among other things, such a feature is expected to inhibit passage ofthrombus from the right atrium to the left atrium, which could causestroke or death. Moreover, allowing flow of blood with low oxygenationfrom right to left would further aggravate CHF. Further, tissue valve130 preferably is configured to close and therefore inhibit flow ineither direction when the pressures at the first and second flared endregions are approximately equal. Preferably, tissue valve 130 is sizedand has dynamic characteristics selected to maintain a pressuredifferential between the left and right atria of 15 mm Hg or less.

To achieve such flow effects, as well as reduce complexity of devicefabrication, tissue valve 130 preferably is a tricuspid valve, as isillustrated in FIGS. 1A-1D, but alternatively may be a bicuspid valve,for example a duckbill valve, or a mitral valve, as described here afterwith respect to FIGS. 8A-8C and 9. For example, as described in greaterdetail below with respect to FIGS. 3A-3E, tissue valve 130 may be formedof a single piece of thinned animal pericardial tissue that is suturedalong at least one edge to form an open-ended conical or ovoid tube, andthen three-dimensionally fixed to assume a normally closed position. Theinlet or bottom (narrower) end of the tube may be coupled, e.g.,sutured, to biocompatible material 120 at or near sinusoidal ring 113,and the sides of the tube optionally may be sutured to struts 111′,111″, and 111′″, as illustrated in FIG. 1D (strut 111′ not shown in FIG.1D). In one embodiment, the bottom end of the tube is sutured tobiocompatible material 120 along substantially straight line 121 that isapproximately 2-3 mm to the right of the narrowest portion of neckregion 104. Without wishing to be bound by theory, it is believed thatsuch a location for line 121 may be sufficiently large as to inhibittissue from atrial septum 210 from growing into tissue valve 130. Inanother embodiment (not illustrated), the bottom end of tissue valve 130is secured, e.g., sutured to biocompatible material 120 along a curvethat follows the shape of sinusoidal ring 113. During use, the outlet orupper (wider) end of the tube may open and close based on the pressuredifferential between the inlet and outlet ends, that is, between theleft and right atria when implanted. Other suitable valve configurationsmay include bicuspid valves, duckbill valves, sleeve (windsock) valves,flap valves, and the like.

As noted above, hourglass-shaped device 100 preferably is configured forimplantation through the fossa ovalis of the atrial septum, particularlythrough the middle of the fossa ovalis. As known to those skilled in theart, the fossa ovalis is a thinned portion of the atrial septum causedduring fetal development of the heart, which appears as an indent in theright side of the atrial septum and is surrounded by a thicker portionof the atrial septum. While the atrial septum itself may be severalmillimeters thick and muscular, the fossa ovalis may be onlyapproximately one millimeter thick, and is formed primarily of fibroustissue. Advantageously, because the fossa ovalis comprises predominantlyfibrous tissue, that region of the atrial septum is not expected toundergo significant tension or contraction during the cardiac cycle, andthus should not impose significant radial stresses on stent 110 thatcould lead to stress-induce cracking. In addition, the composition ofthe fossa ovalis as primarily fibrous tissue is expected to avoidexcessive endothelialization after implantation.

In some embodiments of the present invention, hourglass-shaped device100 is asymmetrically shaped to take advantage of the natural featuresof atrial septum 210 near the fossa ovalis, and to provide suitable flowcharacteristics. FIG. 2A illustrates a plan view of the right atrialside of the atrial septum 210, including an implantation site 201through the fossa ovalis 212. Preferably, the implantation site 201 isthrough the middle of the fossa ovalis 212, so that the device may beimplanted at a spaced distance from the surrounding limbus 214, inferiorvena cava (IVC) 216, and atrial wall 210. For example, as illustrated inFIG. 2B, first flared end region 102 is configured to be implanted inright atrium 204 and may be tapered so as to have a more cylindricalshape than does second flared end region 106, which is configured to beimplanted in left atrium 202. The more cylindrical shape of first flaredend region 102 may enhance opening and closing of tissue valve 130,while reducing risk of the tissue valve falling back towards stent 110;may increase the proportion of tissue valve 130 that moves during eachopen-close cycle, and thus inhibit tissue growth on the valve; and mayreduce or inhibit contact between first flared end region 102 and thelimbus 214 of the fossa ovalis 212, that is, between first flared endregion 102 and the prominent margin of the fossa ovalis, while stillanchoring device 100 across atrial septum 210. The more cylindricalshape of first flared end region 102 further may reduce or inhibitcontact between first flared end region 102 and the right atrial wall,as well as the ridge 218 separating the coronary sinus from the inferiorvena cava (IVC) (shown in FIG. 2A but not FIG. 2B). Additionally, insome embodiments the first flared end region 102 substantially does notextend beyond the indent of the fossa ovalis in the right atrium, andtherefore substantially does not restrict blood flow from the IVC 216.

In accordance with one aspect of the invention, device 100 preferably isconfigured so as to avoid imposing significant mechanical forces onatrial septum 210 or atria 202, 204, allowing the septum to naturallydeform as the heart beats. For example, muscular areas of septum 210 maychange by over 20% between systole and diastole. It is believed that anysignificant mechanical constraints on the motion of atrial septum 210 insuch areas would lead to the development of relatively large forcesacting on the septum and/or on atrial tissue that contacts device 100,which potentially would otherwise cause the tissue to have aninflammatory response and hyperplasia, and possibly cause device 100 toeventually lose patency. However, by configuring device 100 so that neckregion may be implanted entirely or predominantly in the fibrous tissueof the fossa ovalis 212, the hourglass shape of device 100 is expectedto be sufficiently stable so as to be retained in the septum, whilereducing mechanical loads on the surrounding atrial septum tissue 210.As noted elsewhere herein, tissue ingrowth from atrial septum 210 inregions 230 may further enhance binding of device 100 to the septum.

Also, for example, as illustrated in FIG. 2B, neck region 104 of device100 is significantly narrower than flared end regions 102, 106,facilitating device 100 to “self-locate” in a puncture through atrialseptum 210, particularly when implanted through the fossa ovalis. Insome embodiments, neck region 104 may have a diameter suitable forimplantation in the fossa ovalis, e.g., that is smaller than the fossaovalis, and that also is selected to inhibit blood flow rates exceedinga predetermined threshold. For example, the smallest diameter of neck104 may be between about 3 and 6 mm, e.g., between about 4.5 mm and 5.5mm, preferably between about 4.5 mm and 5.5 mm. For example, it isbelieved that diameters of less than about 4.5 mm may in somecircumstances not allow sufficient blood flow through the device todecompress the left atrium, and may reduce long-term patency of device100, while diameters of greater than about 5.5 mm may allow too muchblood flow. For example, flow rates of greater than 1.2 liters/minute,or even greater than 1.0 liters/minute are believed to potentially leadto remodeling of the right atrium. Preferably, the effective diameter atthe narrowest point in device 100, i.e., the narrowest diameter providedby the combination of neck 104 and biocompatible material 120 is about4.5 mm to 4.8 mm. Such a diameter range is expected to provide a flowrate of about 0.80 liters/minute or less following ingrowth of septaltissue, which may anchor device 100 in place, and which may result in anoverall diameter reduction of about 1.0 mm over time.

In some embodiments, the length of first flared end region 102 also maybe selected to protrude into the right atrium by a distance selected toinhibit tissue ingrowth that may otherwise interfere with the operationof tissue valve 130. For example, distance R between the narrowestportion of neck region 104 and the end of first flared region 102 may beapproximately 5.0 to 9.0 mm, for example about 5.5 to about 7.5 mm, orabout 6 mm, so as not to significantly protrude above the limbus offossa ovalis 212. Second flared end region 106 preferably does notsignificantly engage the left side of atrial septum 210, and distance Lmay be between 2.0 and 6.0 mm, for example about 2.5 to 7 mm, or about3.0 mm. It is believed that configuring first and second flared endregions 102, 106 so as to extend by as short a distance as possible intothe right and left atria, respectively, while still maintainingsatisfactory flow characteristics and stabilization in atrial septum210, may reduce blockage of flow from the inferior vena cava (IVC) inthe right atrium and from the pulmonary veins in the left atrium. In oneillustrative embodiment, distance R is about 6.0 mm and distance L isabout 3.0 mm. In some embodiments, the overall dimensions of device 100may be 10-20 mm long (L+R, in FIG. 2B), e.g., about 12-18 mm, e.g.,about 14-16 mm, e.g., about 15 mm.

The diameters of the first and second flared end regions further may beselected to stabilize device 100 in the puncture through atrial septum210, e.g., in the puncture through fossa ovalis 212. For example, firstflared end region 102 may have a diameter of 10-15 mm at its widestpoint, e.g., about 9.0-13 mm; and second flared end region 106 may havea diameter of 10-20 mm at its widest point, e.g., about 13-15 mm. Thelargest diameter of first flared end region 102 may be selected so as toavoid mechanically loading the limbus of the fossa ovalis 212, whichmight otherwise cause inflammation. The largest diameter of secondflared end region 106 may be selected so as to provide a sufficientangle between first and second flared end regions 102, 106 to stabilizedevice 100 in the atrial septum, while limiting the extent to whichsecond flared end region 106 protrudes into the left atrium (e.g.,inhibiting interference with flow from the pulmonary veins), andproviding sufficient blood flow from the left atrium through neck region104. In one embodiment, the angle between the first and second flaredend regions is about 50-90 degrees, e.g., about 60 to 80 degrees, e.g.,about 70 degrees. Such an angle may stabilize device 100 across thefossa ovalis, while inhibiting excessive contact between the device andthe atrial septum. Such excessive contact might cause inflammationbecause of the expansion and contraction of the atrial septum during thecardiac cycle, particularly between diastole and systole. In oneembodiment, the first flared end region subtends an angle ofapproximately 80 degrees, that is, the steepest part of the outersurface of the first flared end region is at an angle of approximately40 degrees relative to a central longitudinal axis of the device. Thesecond flared end region may subtend an angle of approximately 75degrees, that is, the steepest part of the outer surface of the secondflared end region is at an angle of approximately 37.5 degrees relativeto the central longitudinal axis of the device.

Tissue valve 130 is preferably configured such that when closed,leaflets 131 define approximately straight lines resulting from tensionexerted by stent 110 across valve opening 132, as illustrated in FIG.1C. Additionally, the transition between tissue valve 130 andbiocompatible material 120 preferably is smooth, so as to reduceturbulence and the possibility of flow stagnation, which would increasecoagulation and the possibility of blockage and excessive tissueingrowth. As pressure differentials develop across tissue valve 130(e.g., between the left and right atria), blood flow preferably followsa vector that is substantially perpendicular to the tension forcesexerted by stent 110, and as such, the equilibrium of forces isdisrupted and leaflets 131 start to open. As the leaflets open, thedirection of tension forces exerted by stent 110 change, enabling anequilibrium of forces and support of continuous flow. An equilibriumposition for each pressure differential is controlled by the geometry oftissue valve 130 and the elastic behavior of stent 110. When a negativepressure differential (right atrial pressure greater than left atrialpressure) develops, valve leaflets 131 are coapt, closing the tissuevalve and the prevention of right to left backflow.

When device 100 is implanted across the atrial septum, as illustrated inFIG. 2B, left atrial pressures may be regulated in patients havingcongestive heart failure (CHF). For example, device 100 may reducepressure in the left atrium by about 2-5 mmHg immediately followingimplantation. Such a pressure reduction may lead to a long-term benefitin the patient, because a process then begins by which the lowered leftatrial pressure reduces the transpulmonary gradient, which reduces thepulmonary artery pressure. However, the right atrial pressure is notsignificantly increased because the right atrium has a relatively highcompliance. Furthermore, the pulmonary capillaries may self-regulate toaccept high blood volume if needed, without increasing pressure. Whenthe left atrial pressure is high, the pulmonary capillaries constrict tomaintain the transpulmonary gradient, but as the left atrial pressuredecreases, and there is more blood coming from the right atrium, thereare actually higher flow rates at lower pressures passing through thepulmonary circulation. After a period of between a few hours and a weekfollowing implantation of device 100, the pulmonary circulation has beenobserved to function at lower pressures, while the systemic circulationmaintains higher pressures and thus adequate perfusion. The resultinglower pulmonary pressures, and lower left ventricle end diastolicpressure (LVEDP) decrease the after load by working at lower pressures,resulting in less oxygen demand and less resistance to flow. Such smalldecreases in afterload may dramatically increase the cardiac output (CO)in heart failure, resulting in increased ejection fraction (EF).Moreover, because of the release in the afterload and in the pressuresof the pulmonary circulation, the right atrial pressure decreases overtime as well. Following myocardial infarction, the effect is even morepronounced, because the period after the infarction is very importantfor the remodeling of the heart. Specifically, when the heart remodelsat lower pressures, the outcome is better.

In the region of contact between device 100 and atrial septum 210,preferably there is limited tissue growth. The connective tissue ofatrial septum 210 is non-living material, so substantially no nourishingof cells occurs between the septum and device 100. However, localstagnation in flow may lead to limited cell accumulation and tissuegrowth where device 100 contacts atrial septum 210, for example inregions designated 230 in FIG. 2B. Such tissue growth in regions 230 mayanchor device 210 across atrial septum 210. Additional, such tissuegrowth may cause the flow between the external surface of device 100 andatrial septum 210 to become smoother and more continuous, thus reducingor inhibiting further cell accumulation and tissue growth in regions230. As noted above, first flared end region 102 of stent 110, e.g.,between the line along which tissue valve 130 is coupled tobiocompatible material 120 and first sinusoidal ring 112 preferably isbare metal. This configuration is expected to inhibit formation ofstagnation points in blood flow in right atrium 204, that otherwise maylead to tissue growth on the external surfaces of leaflets 131 of tissuevalve 130.

A method 300 of making device 100 illustrated in FIGS. 1A-1D and FIG. 2Bwill now be described with respect to FIGS. 3A-3E.

First, a tube of shape-memory material, e.g., a shape-memory metal suchas nickel titanium (NiTi), also known as NITINOL, is provided (step 301of FIG. 3A). Other suitable materials known in the art of deformablestents for percutaneous implantation may alternatively be used, e.g.,other shape memory alloys, polymers, and the like. In one embodiment,the tube has a thickness of 0.25 mm.

Then, the tube is laser-cut to define a plurality of sinusoidal ringsconnected by longitudinally extending struts (step 302). For example,struts 111 and sinusoidal rings 112-116 illustrated in FIG. 1A may bedefined using laser cutting a single tube of shape-memory metal, andthus may form an integral piece of unitary construction. Alternatively,struts 111 and sinusoidal rings 112-116 may be separately defined fromdifferent pieces of shape-memory metal and subsequently coupledtogether.

Referring again to FIG. 3A, the laser-cut tube then is expanded on amandrel to define first and second flared end regions and a necktherebetween, e.g., to define first end region 102, second end region106, and neck region 104 as illustrated in FIG. 1A; the expanded tubethen may be heated to set the shape of stent 110 (step 303). In oneexample, the tube is formed of NITINOL, shaped using a shape mandrel,and placed into an oven for 11 minutes at 530 C to set the shape.Optionally, the stent thus defined also may be electropolished to reducethrombogenicity, or otherwise suitably treated. Such electropolishingmay alternatively be performed at a different time, e.g., before shapingusing the mandrel.

As shown in FIG. 3A, the neck and second flared end region of the stentthen may be coated with a biocompatible material (step 304). Examples ofsuitable biocompatible materials include expandedpolytetrafluoroethylene (ePTFE), polyurethane, DACRON (polyethyleneterephthalate), silicone, polycarbonate urethane, and animal pericardialtissue, e.g., from an equine, bovine, or porcine source. In oneembodiment, the stent is coated with the biocompatible material bycovering the inner surface of the stent with a first sheet of ePTFE, andcovering the outer surface of the stent with a second sheet of ePTFE.The first and second sheets first may be temporarily secured together tofacilitate the general arrangement, e.g., using an adhesive, suture, orweld, and then may be securely bonded together using sintering to form astrong, smooth, substantially continuous coating that covers the innerand outer surfaces of the stent. Portions of the coating then may beremoved as desired from selected portions of the stent, for exampleusing laser-cutting or mechanical cutting. For example, as shown in FIG.1A, biocompatible material 120 may cover stent 110 between sinusoidalring 113 and sinusoidal ring 116, i.e., may cover neck region 104 andsecond flared end region 106, but may be removed between sinusoidal ring113 and sinusoidal ring 112, i.e., may be removed from (or not appliedto) first flared end region 102.

The biocompatible material facilitates funneling of blood from the leftatrium to the right atrium by facilitating the formation of a pressuregradient across tissue valve 130, as well as providing a substantiallysmooth hemodynamic profile on both the inner and outer surfaces ofdevice 100. Advantageously, this configuration is expected to inhibitthe formation of eddy currents that otherwise may cause emboli to form,and facilitates smooth attachment of the device to the atrial septum,e.g., fossa ovalis. Biocompatible material 120 preferably is configuredso as to direct blood flow from the left atrium, through neck region 104and toward tissue valve leaflets 131. Biocompatible material 120preferably also is configured so as to inhibit tissue growth from atrialseptum 210 and surrounding tissue into device 100 and particularlytoward tissue valve leaflets 131. In some embodiments, the biocompatiblematerial 120 has a porosity that is preselected to allow limited cellgrowth on its surface; the cells that grow on such a surface preferablyare endothelial cells that are exposed to blood and inhibit blood fromcoagulating on the biocompatible material. After such cells grow on thebiocompatible material 120, the material preferably is substantiallyinert and thus not rejected by the body. Optionally, the biocompatiblematerial may be impregnated with a second material that facilitatestissue ingrowth, e.g., carbon. Such impregnation may be performed beforeor after applying the biocompatible material to the stent.

Then, as shown in FIG. 3A, a valve having two or more leaflets, such asa tricuspid, bicuspid, or duckbill valve, or any other suitable valve,is formed by folding and suturing a sheet of thinned animal pericardialtissue, e.g., equine, bovine, or porcine material (step 305). FIGS.3B-3E illustrate plan views of exemplary sheets of animal pericardialtissue that may be used to form tissue valves. Specifically, FIG. 3Billustrates an approximately semicircular sheet 310 of tissue for use inpreparing a tricuspid tissue valve. Although the sheet 310 may be anysuitable dimensions, in the illustrated embodiment the sheet has a widthof 10-16 mm, a length of 6-8 mm. The opposing edges may be at an anglebetween 0-70 degrees relative to one another so that when the sheet isfolded and those edges are secured, e.g., sutured together, sheet 310forms a generally funnel-like shape having approximately the same angleas the first flared end region to which it is to be secured. FIG. 3Cillustrates an embodiment similar to that of FIG. 3B, but in which sheet320 also includes wings 321 providing additional tissue material inregions along the suture line that may be subjected to high stresses, aswell as a curved top contour 322 that provides an extended region forcoaptation between the leaflets when the valve is closed. Wings may beapproximately 2-5 mm long, and extend 0.5-1.5 mm beyond the lateraledges of sheet 320. FIG. 3D illustrates an embodiment similar to that ofFIG. 3C, e.g., that includes wings 331 that may be of similar dimensionas wings 321, but in which sheet 330 lacks a curved top contour. Sutures332 are shown in FIG. 3D. FIG. 3E illustrates a sheet 340 of tissuesuitable for use in preparing a bicuspid tissue valve, that has agenerally rectangular shape, for example having a width of 14-15 mm anda length of 6.0-7.0 mm. Other dimensions may suitably be used. Forexample, the tissue sheet may have a flattened length of no greater than18 mm, for example, a length of 10-16 mm, or 12-14 mm, or 14-18 mm, andmay be folded and sutured to define two or more leaflets each having alength of, for example, 9 mm or less, or 8 mm or less, or 7 mm or less,or 6 mm or less, or even 5 mm or less, e.g., 5-8 mm. The tissue sheetmay have a flattened height no greater than 10 mm, for example, a heightof 2-10 mm, or 4-10 mm, or 4-8 mm, or 6-8 mm, or 4-6 mm. The tissuesheet may have a flattened area of no greater than 150 square mm, forexample, 60-150 square mm, or 80-120 square mm, or 100-140 square mm, or60-100 square mm. In some exemplary embodiments, the sheet of tissue mayhave a generally trapezoidal or “fan” shape, so that when opposing edgesare brought together and sutured together, the sheet has a general“funnel” shape, with a wide opening along the outlet or upper edge and anarrow opening along the inlet or lower edge. Note that other suitablemethods of securing opposing edges of the sheet alternatively may beused, e.g., adhesive, welding, and the like.

The tissue may have a thickness, for example, of between 0.050 mm and0.50 mm, for example, about 0.10 mm and 0.20 mm. Typically, harvestedbovine pericardial tissue has a thickness between about 0.3 mm and 0.5mm, which as is known in the art is a suitable thickness for high-stressapplications such as construction of aortic valves. However, for use inthe device of the present invention, it may be preferable to thin thepericardial tissue. For example, the stresses to which the valveleaflets are exposed in a device constructed in accordance with thepresent invention may be a small fraction (e.g., 1/25th) of the stressesin an aortic valve application, because of the relatively large surfacearea of the leaflets and the relatively low pressure gradients acrossthe device. For this reason, thinned pericardial tissue may be used,enabling construction of a more compliant valve that may be readilyfixed in a normally closed position but that opens under relatively lowpressure gradients. Additionally, the use of thinner leaflets isexpected to permit the overall profile of the device to be reduced inwhen the device is compressed to the contracted delivery state, therebyenabling its use in a wider range of patients.

For example, harvested pericardial tissue typically includes threelayers: the smooth and thin mesothelial layer, the inner looseconnective tissue, and the outer dense fibrous tissue. The pericardialtissue may be thinned by delaminating and removing the dense fibroustissue, and using a sheet of the remaining mesothelial and looseconnective layers, which may have a thickness of 0.10 mm to 0.20 mm, toconstruct the tissue valve. The dense fibrous tissue may be mechanicallyremoved, for example using a dermatome, grabbing tool, or by hand, andany remaining fibers trimmed.

The animal pericardial tissue then may be three-dimensionally shaped ona mandrel to define a tissue valve having valve leaflets that arenormally in a closed position, and then fixed in that position usingglutaraldehyde or other suitable substance (step 306). Excessglutaraldehyde may be removed using an anticalcification treatment, forexample to inhibit the formation of calcium deposits on the tissuevalve.

The outlet or upper (wider) portion of the tissue valve then may besecured, e.g., sutured, to the first flared end region, and the inlet orlower (narrower) portion of the tissue valve secured, e.g., sutured tothe biocompatible polymer at the neck region (step 307). For example, asillustrated in FIGS. 1A-1D, the lower portion of tissue valve 130 may besecured using sutures to biocompatible material 120 at or nearsinusoidal ring 113 (for example, along a line 121 approximately 2-3 mmto the right of the narrowest portion of neck region 104), and also maybe sutured to elongated struts 111′, 111″, and 111′″ so as to define atricuspid valve having leaflets 131. Other suitable methods of securingthe tissue valve to stent 110 and to biocompatible material 120 mayalternatively be used. Preferably, tissue valve 130 is secured to device100 such that, when implanted, the tissue valve is disposedsubstantially only in the right atrium. Such a configuration mayfacilitate flushing of the external surfaces of leaflets 131 with bloodentering the right atrium. By comparison, it is believed that ifleaflets 131 were instead disposed within neck region 104 or secondflared end region 106, they might inhibit blood flow and/or graduallylose patency over time as a result of tissue ingrowth caused by thestagnation of blood around the leaflets.

A method 400 of using device 100 illustrated in FIGS. 1A-1D to reduceleft atrial pressure in a subject, for example, a human having CHF, willnow be described with reference to FIG. 4. Some of the steps of method400 may be further elaborated by referring to FIGS. 5A-5D.

First, an hourglass-shaped device having a plurality of sinusoidal ringsconnected by longitudinally extending struts that define first andsecond flared end regions and a neck disposed therebetween, as well as atissue valve coupled to the first flared end region, is provided (step401). Such a device may be provided, for example, using method 300described above with respect to FIGS. 3A-3E.

Then, the device is collapsed radially to a contracted delivery state,and loaded into a loading tube (step 402). For example, as illustratedin FIGS. 5A-5B, device 100 may be loaded into loading tube 510 usingpusher 520 having “star”-shaped end 521. Loading tube 510 includestapered loading end 511, which facilitates radial compression of device100 into lumen 512 having a suitable internal diameter. Once device 100is loaded into lumen 512, pusher 520 is retracted. Preferably, device100 is loaded into loading tube 510 shortly before implantation, so asto avoid unnecessarily compressing device 100 or re-setting of theclosed shape of leaflets 132, which may interfere with later deploymentor operation of the device. In some embodiments, loading tube 510 has adiameter of 16 F or less, or 14 F or less, or 10 F or less, or 6 F orless, e.g., about 5 F, and device 100 has a crimped diameter of 16 F orless, or 14 F or less, or 10 F or less, or 6 F or less, e.g., about 5 F.In one illustrative embodiment, loading tube has a diameter of 15 F anddevice 100 has a crimped diameter of 14 F.

Referring again to FIG. 4, the device then is implanted, first byidentifying the fossa ovalis of the heart septum, across which device100 is to be deployed (step 403). Specifically, a BROCKENBROUGH needlemay be percutaneously introduced into the right atrium via the subject'svenous vasculature, for example, via the femoral artery. Then, underfluoroscopic or echocardiographic visualization, the needle is pressedagainst the fossa ovalis, at a pressure insufficient to puncture thefossa ovalis. As illustrated in FIG. 5C, the pressure from needle 530causes “tenting” of fossa ovalis 541, i.e., causes the fossa ovalis tostretch into the left atrium. Other portions of atrial septum 540 arethick and muscular, and so do not stretch to the same extent as thefossa ovalis. Thus, by visualizing the extent to which differentportions of the atrial septum 540 tents under pressure from needle 530,fossa ovalis 541 may be identified, and in particular the centralportion of fossa ovalis 541 may be located.

Referring again to FIG. 4, the fossa ovalis (particularly its centralregion) may be punctured with the BROCKENBROUGH needle, and a guidewiremay be inserted through the puncture by threading the guidewire throughthe needle and then removing the needle (step 404, not illustrated inFIG. 5). The puncture through the fossa ovalis then may be expanded byadvancing a dilator over the guidewire. Alternatively, a dilator may beadvanced over the BROCKENBROUGH needle, without the need for aguidewire. The dilator is used to further dilate the puncture and asheath then is advanced over the dilator and through the fossa ovalis;the dilator and guidewire or needle then are removed (step 405, notillustrated in FIG. 5). The loading tube, with device 100 disposed in acontracted delivery state therein, then is advanced into the sheath(step 406, not illustrated in FIG. 5).

The device then is advanced out of the loading tube and into the sheathusing a pusher, and then partially advanced out of the sheath, such thatthe second flared end of the device protrudes out of the sheath and intothe left atrium, and expands to its deployed state (step 407). Forexample, as illustrated in FIG. 5D, pusher 550 may be used to partiallyadvance device 100 out of sheath 512 and into left atrium 502, whichcauses the second flared end region to expand in the left atrium. Thepusher may be configured such that it cannot advance the device 100completely out of the sheath, but instead may only push out the side ofthe device to be disposed in the left atrium, that is, the second flaredend region. After the pusher advances the second flared end region outof the sheath, the pusher may be mechanically locked from advancing thedevice out any further. For example, an expanded region may be disposedon the end of the pusher proximal to the physician that abuts the sheathand prevents further advancement of the pusher after the second flaredend region is advanced out of the sheath. The device then may be fullydeployed by pulling the sheath back, causing the second flared endregion of the device to engage the left side of the atrial septum. Sucha feature may prevent accidentally deploying the entire device in theleft atrium.

The sheath then is retracted, causing the second flared end region toflank the left side of the atrial septum and the neck of the device tolodge in the puncture through the fossa ovalis, and allowing expansionof the first flared end of the device into the right atrium (step 408,see also FIG. 2B). Any remaining components of the delivery system thenmay be removed, e.g., sheath, and loading tube (step 409). Oncepositioned in the fossa ovalis, the device shunts blood from the leftatrium to the right atrium when the left atrial pressure exceeds theright atrial pressure (step 410), thus facilitating treatment and/or theamelioration of symptoms associated with CHF.

The performance characteristics of device 100 were characterized usingcomputational fluid dynamic modeling. FIG. 6A is a cross-sectional imageof fluid flow through device 100 in the open configuration, in whichintensity indicates fluid velocity through the device. As can be seen inFIG. 6A, there are substantially no points of stagnation or turbulencein the blood flow. The maximum shear stresses within device 100 werecalculated to be about 50-60 Pascal, which is significantly lower thanvalues that may lead to thrombus formation, which are above 150 Pascal.

The performance of device 100 was also characterized using hemodynamictesting. FIG. 6B is a plot of the flow rate through device 100 as afunction of the pressure differential between the left and right atria,for devices having inner diameters of 3.5 mm (trace 610), 4.2 mm (trace620), 4.8 mm (trace 630), and 5.2 mm (trace 640). At a pressuredifferential of 10 mm Hg, it can be seen that the flow rate of the 3.5mm device was 670 ml/minute; the flow rate of the 4.2 mm device was 1055ml/minute; the flow rate of the 4.8 mm device was 1400 ml/minute; andthe flow rate of the 5.2 mm device was 1860 ml/minute. Based on thesemeasurements, it is believed that devices having inner diameters of 4.5mm to 4.8 mm may provide suitable flow parameters over time, whenimplanted, because ingrowth of septal tissue over the first 6 monthsfollowing implantation may reduce the inner diameter to about 3.5 to 3.8mm, thus reducing the flow rate to below about 800 ml/minute. At steadystate, such a flow rate may reduce the left atrial pressure by 5 mmHg,to around 10-15 mmHg, and may reduce the pressure differential betweenthe left and right atria to about 4-6 mmHg.

Additionally, device 100 was subjected to an accelerated wear andfatigue test for up to 100 million cycles to simulate and predictfatigue durability, and was observed to perform satisfactorily.

The devices and methods described herein may be used to regulate leftatrial pressures in patients having a variety of disorders, includingcongestive heart failure (CHF), as well as other disorders such aspatent foramen ovale (PFO), or atrial septal defect (ASD). The devicesand methods also may be used to reduce symptoms and complicationsassociated with such disorders, including myocardial infarction. It isbelieved that patients receiving the device may benefit from betterexercise tolerance, less incidence of hospitalization due to acuteepisodes of heart failure, and reduced mortality rates.

The devices and methods described herein further may be used tonon-invasively determine the pressure in the left atrium, and thus toassess the efficacy of the device and/or of any medications beingadministered to the patient. Specifically, with respect to FIG. 7,method 700 includes imaging an implanted hourglass-shaped device, e.g.,device 100 described above with respect to FIGS. 1A-1D (step 701). Suchimaging may be ultrasonic, e.g., cardioechographic, or may befluoroscopic. Using such imaging, the time duration of the opening oftissue valve 130 may be measured (step 702). Based on the measured timeduration, the flow of blood through the valve may be calculated (step703). The left atrial pressure then may be calculated based on thecalculated flow, for example, based on a curve such as shown in FIG. 6B(step 704). Based on the calculated left atrial pressure, the efficacyof the valve and/or of any medication may be assessed (step 705). Aphysician may adjust the medication and/or may prescribe a new treatmentplan based on the assessed efficacy of the valve and/or the medication.

Some alternative embodiments of device 100 described above with respectto FIGS. 1A-1D are now described. In particular, tissue valves otherthan tricuspid valve 130 illustrated above with respect to FIGS. 1A-1Dmay be employed with device 100. For example, device 800 illustrated inFIGS. 8A-8C includes hourglass-shaped stent 110, which may besubstantially the same as stent 110 described above, biocompatiblematerial 120, and duckbill tissue valve 830. Like device 100, device 800has three general regions: first flared or funnel-shaped end region 102configured to flank the right side of the atrial septum, second flaredor funnel-shaped end region 106 configured to flank the left side of theatrial septum, and neck region 104 disposed between the first and secondflared end regions and configured to lodge in a puncture formed in theatrial septum, preferably in the fossa ovalis. Stent 110 includesplurality of sinusoidal rings 112-116 interconnected by longitudinallyextending struts 111, which may be laser cut from a tube of shape memorymetal. Neck region 104 and second flared end region 106 may be coveredwith biocompatible material 120, e.g., in the region extendingapproximately from sinusoidal ring 113 to sinusoidal ring 116.

Duckbill tissue valve 830 is coupled to stent 110 in first flared endregion 102. Preferably, tissue valve 830 opens at a pressure of lessthan 1 mmHg, closes at a pressure gradient of 0 mmHg, and remains closedat relatively high back pressures, for example at back pressures of atleast 40 mmHg. Like tissue valve 130, tissue valve 830 may be formedusing any natural or synthetic biocompatible material, including but notlimited to pericardial tissue, e.g., thinned and fixed bovine, equine,or porcine pericardial tissue. As shown in FIG. 8B, the outlet ofduckbill tissue valve 830 is coupled, e.g., sutured, to first and secondlongitudinally extending struts 111′, 111″ in the region extendingbetween first (uppermost) sinusoidal ring 112 and second sinusoidal ring113. Referring again to FIG. 8A, the inlet to tissue valve 830 also iscoupled, e.g., sutured, to the upper edge of the biocompatible material120 along line 121, at or near sinusoidal ring 113, so as to provide asmooth profile.

FIGS. 8A and 8B illustrate device 800 when duckbill tissue valve 830 isin an open configuration, in which leaflets 931 are in an open positionto permit flow. FIG. 8C illustrates device 800 when tissue valve 830 isin a closed configuration, in which leaflets 831 are in a closedposition to inhibit flow, in which position they preferably form asubstantially straight line. Device 800 preferably is configured so asto provide flow characteristics similar to those described above fordevice 100.

Referring now to FIG. 9, alternative device of the present invention isdescribed. Device 900 has first and second flared end regions 902, 906,with neck region 904 disposed therebetween. Device 900 includeshourglass-shaped stent 910, biocompatible material 920, and tissue valve930 and further comprises three general regions as described for theforegoing embodiments: first flared or funnel-shaped end region 902configured to flank the right side of the atrial septum, second flaredor funnel-shaped end region 906 configured to flank the left side of theatrial septum, and neck region 904 disposed between the first and secondflared end regions and configured to lodge in a puncture formed in theatrial septum, preferably in the fossa ovalis. Like the devicesdescribed above, stent 910 includes plurality of sinusoidal rings 912interconnected by longitudinally extending struts 911, which may belaser cut from a tube of shape memory metal. However, as compared todevices 100 and 800 described further above, sinusoidal rings 912 do notextend into first flared end region 902. Instead, the outlet end oftissue valve 930 is coupled to longitudinally extending struts 911′ and911″. Neck region 904 and second flared end region 906 may be coveredwith biocompatible material 920.

Duckbill tissue valve 930 is coupled to stent 910 in first flared endregion 902. Specifically, the outlet of tissue valve 930 is coupled,e.g., sutured, to first and second longitudinally extending struts 911′,911″ in the region extending between the first (uppermost) sinusoidalring 912 and the distal ends of struts 911′, 911″. The inlet end oftissue valve 930 also is coupled, e.g., sutured, to the upper edge ofbiocompatible material 920 at or near first (uppermost) sinusoidal ring912, so as to provide a smooth profile. Device 900 is preferablyconfigured so as to provide flow characteristics similar to thosedescribed above for device 100.

EXAMPLE

An exemplary device 800 such as described above with respect to FIGS.8A-8C was implanted into four sheep with induced chronic heart failure(V1-V4), while four sheep with induced chronic heart failure did notreceive the device, and were used as a control (C1-C4). An additionalcontrol animal was subjected to only a partial heart failure protocol,and did not receive the device (S1).

Chronic heart failure was induced in animals C1-C4 and V1-V4, who wereless than 1 year of age and weighed between 70 and 120 pounds, by firstanesthetizing the animals via a venous catheter positioned in aperipheral vessel, i.e., the ear. The animals were given an opiate orsynthetic opiate (e.g., morphine or butorphanol) intravenously at 0.25to 0.5 mg/kg, as well as telazol at 0.3 mg/kg, through the venouscatheter, and anesthetized by intravenous etomidate. Anesthesia wasmaintained with 1.5% isoflurane delivered in 100% O₂, via a trachealtube. The animals were placed on a fluoroscope table in left lateralrecumbence, and a gastric tube (about 7 F) was inserted into the rumento serve as a vent.

An introducer was then positioned within the carotid artery via cut downand modified Seldinger technique. A 6 F or 7 F Judkins left 4.5 catheterwas advanced through the introducer into the left circumflex coronaryartery (LCxA) under fluoroscopic guidance, and about 60,000 polystyrenemicrospheres of about 90 μm diameter were injected into the LCxA toinduce embolization to induce myocardial infarction followed by chronicheart failure. The arterial and skin incisions then were closed, and theanimals were administered about 500 mg of cephalexein p.o. bid for twodays, as well as a synthetic opiate prn, specifically buprenorphineadministered intramuscularly at about 0.03 to 0.05 mg/kg, once duringrecovery and following the anesthesia. Animals observed to havearrhythmia following or during the microsphere injection were alsoadministered lidocaine following embolization, at about 2 to 4 mg/kg viaintravenous bolus, followed by constant infusion at about 20 to 80μf/kf/minute.

This procedure was repeated one week following the first procedure inanimals V1-V4 and C1-C4. This model of induced chronic heart failure hasabout a 100% fatality rate at 12 weeks, and as discussed below each ofthe control animals died before the end of the 12 week study. Theprocedure was performed a single time in animal S1, and as discussedbelow this animal survived the 12 week study but deteriorated over thecourse of the study.

Device 800 was implanted into four animals V1-V4. Fluid filled catheterswere implanted into animals V1-V4 and C1-C4, approximately seven daysafter the second embolization procedure. Fluid filled catheters were notimplanted into animal S1. The implanted device 800 had an overall lengthof 15 mm (7 mm on the left atrial side and 8 mm on the right atrialside), a diameter on the left atrial side of 14 mm, a diameter on theright atrial side of 13 mm, an inside neck diameter of 5.3 mm, and anangle between the left and right atrial sides of the device of 70degrees. The fluid filled catheters were implanted in the inferior venacava (IVC), superior vena cava (SVC), pulmonary artery, and left atriumthrough a right mini-thoracotomy under anesthesia, and were configuredto measure oxygen saturations and pressures in the IVC, pulmonaryartery, right atrium, and left atrium. After implantation and throughoutthe study, the animals were each treated daily with aspirin, plavix, andclopidogrel. Their heart rate was periodically monitored.

Two-dimensional M-mode echocardiograms of the left ventricle wereperiodically obtained to document the ejection fraction (EF), as well asthe shortening fraction, calculated as 100(EDD−ESD)/EDD, where EDD isthe end-diastolic dimension (diameter across ventricle at the end ofdiastole) and ESD is the end-systolic dimension (diameter acrossventricle at the end of systole). Echocardiographic studies of theanimals were performed while they were either conscious or under lightchemical restraint with butorphanol, and manually restrained in theright or left decubitis position, using an ultrasound system with a 3.5to 5.0 mHz transducer (Megas ES, model 7038 echocardiography unit). Theechocardiograms were recorded for subsequent analysis. The leftventricle fractional area shortening (FAS), a measure of left ventriclesystolic function, was measured from the short axis view at the level ofthe papillary muscles. Measurements of left ventricle dimensions,thickness of the posterior wall, and intraventricular septum wereobtained and used as an index of left ventricle remodeling. The majorand minor axes of the left ventricle were measured and used to estimateleft ventricle end-diastolic circumferential wall stress.

The clinical conditions of the animals were evaluated by comparingvarious parameters over a twelve-week period, including left atrialpressure, right atrial pressure, pulmonary artery pressure, and ejectionfraction (EF). Parameters such as left and right atrial pressures, leftand right ventricular dimensions, and left and right ventricularfunction were obtained based on the collected data. Data obtained duringthe study are discussed further below with respect to FIGS. 10A-10D andTables 2-15.

During the course of the study, all four of the control animals C1-C4were observed to suffer from high pulmonary artery pressure, high rightatrial pressure, and low ejection fraction, and were immobile. All fourcontrol animals died during the trial, C3 at week 1, C4 at week 3, C1 atweek 6, and C2 at week 9. Animal S1 survived but deteriorated over thecourse of the study.

By comparison, all of the animals V1-V4 into which the device had beenimplanted were observed to have dramatically improved hemodynamicconditions over the course of the study, and appeared healthy andenergetic without signs of congestion by the end of the study. Asdiscussed below with reference to FIGS. 10A-10D, device 800 was observedto reduce left atrial pressure in the implanted animals by about 5 mmHg,with an increase in cardiac output, and preservation of right atrialpressure and pulmonary artery pressure. Left ventricle parameters wereobserved to be substantially improved in the implanted animals ascompared to the control animals, and right ventricle and pulmonaryartery pressure were also observed to be normal in the implantedanimals.

Three of the four implanted animals, V1, V3, and V4 survived the twelveweek study. One of the implanted animals, V2, died at week 10 of anon-heart failure cause. Specifically, arrhythmia was diagnosed as thecause of death; the animal was observed to have arrhythmia at baseline,and had been defibrillated before implantation Throughout the study,this animal was observed to have good hemodynamic data. At the end ofthe study, the surviving implant animals were observed to respondnormally to doses of dobutamine, indicating significant improvement inthe condition of their heart failure.

FIG. 10A is a plot of the measured left atrial pressure of the controlanimals (C1-C4), and of the implanted animals (V1-V4), along with meanvalues for each (M.C. and M.V., respectively). Data for control animalC3 is not shown, as the animal died in the first week of the study. Themean left atrial pressure for the control animals (M.C.) was observed tosteadily increase over the course of the study, from about 14 mmHg atbaseline to over 27 mmHg when the last control animal (C1) died. Bycomparison, the mean left atrial pressure for the implanted animals(M.V.) was observed to drop from about 15 mmHg at baseline to less than12 mmHg at week one, and to remain below 14 mmHg throughout the study.

FIG. 10B is a plot of the measured right atrial pressure of the controlanimals (C1-C4), and of the implanted animals (V1-V4), along with meanvalues for each (M.C. and M.V., respectively). Data for control animalC3 is not shown. As for the left atrial pressure, the mean right atrialpressure for the control animals (M.C.) was observed to steadilyincrease over the course of the study, from about 5.5 mmHg at baselineto over 12 mmHg when the last control animal (C1) died. By comparison,the mean right atrial pressure for the implanted animals (M.V.) wasobserved to remain relatively steady throughout the study, increasingfrom about 6 mmHg to about 7 mm Hg over the first two weeks of thestudy, and then decreasing again to about 6 mmHg for the rest of thestudy.

FIG. 10C is a plot of the measured ejection fraction of the controlanimals (C1-C4), and of the implanted animals (V1-V4), along with meanvalues for each (M.C. and M.V., respectively). Data for control animalC3 is not shown. The mean ejection fraction for the control animals(M.C.) was observed to steadily decrease over the course of the study,from about 38% at baseline to about 16% when the last control animal(C1) died. By comparison, the mean ejection fraction for the implantedanimals (M.V.) was observed to steadily increase over the course of thestudy, from about 33% at baseline to about 46% at the conclusion of thestudy.

FIG. 10D is a plot of the measured pulmonary artery pressure of thecontrol animals (C1-C4), and of the implanted animals (V1-V4), alongwith mean values for each (M.C. and M.V., respectively). Data forcontrol animal C3 is not shown. The mean pulmonary artery pressure forthe control animals (M.C.) was observed to vary significantly over thecourse of the study, from about 27 mmHg during the first week of thestudy, to about 45 mmHg at week six, then down to 40 mmHg at week eight,and then up to about 47 mmHg at week nine, when the last control animal(C1) died. By comparison, the mean pulmonary artery pressure for theimplanted animals (M.V.) was observed to remain relatively steady,increasing from about 22 mmHg during week one, to about 27 mmHg duringweeks four through nine, and then back down to about 24 mmHg by weektwelve, at the conclusion of the study.

Upon explanation at the end of the study, three of the four implanteddevices were observed to be completely patent and functional. Forexample, FIGS. 11A-11B are photographic images of device 800 uponexplanation from one of the implanted animals, taken from the leftatrial and right atrial sides respectively. A fourth device was observedto be patent up until week 11, using Fick's measurements andechocardiography. At histopathology, no inflammation was observed aroundthe valves, and a thin endothelial layer was observed to have ingrown.For example, FIG. 11C is a microscope image of device 800 uponexplanation from one of the implanted animals, showing approximately 0.2mm of endothelial tissue in the device in the neck region.

Tables 2 through 15 present raw data obtained from the control animalsC1-C4 and S1 and the implanted animals V1-V4, while awake, over thecourse of the 12 week study, including baseline immediately beforeimplantation (Day 0, during which the animals were sedated). The meanvalues for control animals C1-C4 and S1 (M.C.) and the mean values forthe implanted animals V1-V4 (M.V.), with standard deviations, are alsopresented in the tables. Missing data indicates either the death of theanimal or omission to obtain data. Data for animal C3 is not shownbecause the animal died in the first week of the study. Data was notcollected for any animal in week 7 of the study. As noted above, animalS1 was not implanted with pressure and saturation flow monitors, so nodata is shown for that animal for certain measurements.

Table 2 presents the study's results pertaining to right atrial pressure(RAP, mmHg). As can be seen from Table 2, the average RAP for thecontrol animals (C1-C4) increased significantly over the course of thestudy. For example, animal C1 experienced an RAP increase to about 330%of baseline before death, C2 to about 110% of baseline before death, andC4 to about 340% of baseline before death. The increase was relativelysteady during this period. By contrast, the RAP for the implantedanimals (V1-V4) started at a similar value to that of the controlanimals, at an average of 6±2 mmHg at baseline, but did notsignificantly vary over the course of the study. Instead, the averageRAP of the implanted animals remained within about 1-2 mmHg of thebaseline value for the entire study (between a high of 7±1 and a low of5±1). Thus, the inventive device may inhibit increases in the rightatrial pressure in subjects suffering from heart failure, and indeed maymaintain the right atrial pressure at or near a baseline value. This isparticularly noteworthy because, as described elsewhere herein, thedevice may offload a relatively large volume of blood from the leftatrium to the right atrium; however the relatively high compliance ofthe right atrium inhibits such offloading from significantly increasingRAP.

TABLE 2 Right Atrial Pressure (RAP, mmHg) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk. 4Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 3.8 4.3 5.1 4.1 10.811.6 12.1 12.8 12.6 C2 9.2 10.1 10.5 9.8 8.6 9.8 10.3 C4 3.3 5.7 6.111.4 S1 V1 8.9 7.1 8.2 5.6 6.8 5.7 6.1 6.9 7.1 6.5 5.7 6.3 V2 7.4 6.16.7 5.5 5.6 6.0 6.4 7.0 6.5 V3 8.0 7.7 7.7 7.6 6.7 6.0 5.5 5.8 5.4 6.77.2 5.7 V4 0.9 5.2 5.1 4.9 5.7 5.8 3.4 3.8 4.8 5.0 5.6 5.7 M.C. 5 ± 2 7± 2 7 ± 1 8 ± 2 10 ± 1 11 ± 1 11 ± 1 13 13 M.V. 6 ± 2 7 ± 1 7 ± 1 6 ± 1 6 ± 0  6 ± 0  5 ± 1 6 ± 1 6 ± 1 6 ± 1 6 ± 1 6 ± 0

Table 3 presents the study's results pertaining to left atrial pressure(LAP, mmHg). As can be seen from Table 3, the average LAP of the controlanimals started at a similar value at baseline as that of the implantedanimals, 14±1 mmHg for the former and 15±12 mmHg for the latter.However, the LAP of the control animals increased significantly over thecourse of the study. For example, animal C1 had a baseline LAP of 10.6mmHg, and an LAP of 27.3 mmHg at week 9 just before death, about 250% ofbaseline. The LAP increases of the other control animals were smaller,but still significantly larger than that of the implanted animals.Indeed, in each case the LAP of the implanted animals actually decreasedimmediately following implantation. For example, the LAP for animal V1decreased from 15.7 mmHg at baseline to 11.4 mmHg one week followingimplantation, about 73% of baseline. The average LAP for the implantedanimals decreased from 15±2 at baseline to a low of 11±0 at week one,and then gradually increased to about 13±1 at week six (about 87% ofbaseline), where it remained for the remainder of the study.

TABLE 3 Left Atrial Pressure (LAP, mmHg) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk. 4Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 10.6 12.8 15.9 13.6 17.023.5 24.4 26.0 27.3 C2 14.4 15.1 16.3 18.1 18.1 19.7 20.7 C4 16.4 17.718.9 23.7 S1 V1 15.7 11.4 11.3 8.8 9.2 13.4 14.3 15.0 14.9 13.9 15.215.6 V2 19.8 11.7 11.7 12.1 12.3 13.0 14.7 14.2 14.0 V3 14.3 12.1 12.412.7 12.0 11.5 11.6 11.8 11.9 12.4 13.0 12.3 V4 10.3 10.1 11.3 11.4 11.010.2 10.8 11.2 11.7 11.9 12.2 12.1 M.C. 14 ± 1 15 ± 1 17 ± 1 18 ± 3 18 ±0 22 ± 2 23 ± 2 26 27 M.V. 15 ± 2 11 ± 0 12 ± 0 11 ± 1 11 ± 1 12 ± 1 13± 1 13 ± 1 13 ± 1 13 ± 1 13 ± 1 13 ± 1

Table 4 further elaborates the results presented in Table 3, andpresents the calculated change in LAP (ΔLAP, %). As can be seen in Table4, control animals C2 and C4 each died after their LAP increased byabout 44%, while control animal C1 died after its LAP increased by about158%. By comparison, implanted animals V1, V2, and V3 each experiencedsignificant decreases in LAP immediately following implantation, e.g.,by about −27%, −41%, and −15% relative to baseline. The LAP for animalV4 remained near baseline following implantation. The LAP for animal V1slowly increased back to baseline over the course of the study; the LAPfor animal V2 remained significantly below baseline before its death butincreased somewhat; the LAP for animal V3 also remained below baselinethroughout the study but increased somewhat; and the LAP for animal V4fluctuated somewhat above baseline but remained within about 18% ofbaseline. Thus, it can be seen that the inventive device may inhibitincreases in the left atrial pressure in patients suffering from heartfailure. Indeed, the device may actually decrease the left atrialpressure below baseline in patients suffering from heart failure for atime period immediately following implantation, in some embodiments to alevel about 20% below baseline. The left atrial pressure subsequentlymay gradually increase back towards a baseline level over a time periodof weeks or months, as the heart remodels and improves in efficiency. Itis important to note that the control animals died from pulmonary edema,which correlates with LAPs that exceed the “danger zone” of 25 mmHg ormore at which edema occurs.

TABLE 4 Change in Left Atrial Pressure (ΔLAP, %) Day 0 Wk. 1 Wk. 2 Wk. 3Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 +5 +51 +29 +61+122 +131 +145 +158 C2 +21 +14 +26 +26 +37 +44 C4 +8 +15 +44 S1 V1 −27−28 −44 −41 −15 −9 −4 −5 −11 −3 0 V2 −41 −41 −39 −38 −34 −26 −28 −29 V3−15 −13 −11 −16 −20 −19 −17 −16 −13 −9 −13 V4 −2 +10 +10 +7 −1 +5 +8 +13+16 +18 +17 M.C. +11 ± 4 +27 ± 10 +33 ± 5  +44 ± 14 +80 ± 42 +87 ± 35+145 +158 M.V. −21 ± 8 −18 ± 11 −21 ± 13 −22 ± 11 −17 ± 7  −12 ± 7  −10± 8 −9 ± 9 −3 ± 9 +2 ± 8 +1 ± 9

Table 5 presents the study's results pertaining to pulmonary arterypressure (PAP, mmHg). As can be seen in Table 5, the control animalsexperienced significant increases in PAP before death, e.g., about 230%of baseline for animal C1, 217% of baseline for animal C2, and 180% ofbaseline for animal C4. The PAP for the implanted animals also increasedover the course of the study, but in most cases by significantly lessthan that of the control animals, e.g., to about 133% of baseline foranimal V1, about 161% of baseline for animal V2, about 156% of baselinefor animal V3, and about 169% for animal V4. The inventive device thusmay inhibit increases in pulmonary artery pressure in subjects sufferingfrom heart failure, relative to what they may otherwise have experiencedduring heart failure.

TABLE 5 Pulmonary Artery Pressure (PAP, mmHg) Day 0 Wk. 1 Wk. 2 Wk. 3Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 20.8 27.9 28.527.9 28.0 41.7 40.2 48.0 C2 22.3 25.8 29.7 26.9 32.0 43.5 48.4 C4 20.128.4 31.2 36.1 S1 V1 18.6 21.2 20.7 27.1 30.2 28.4 29.0 29.8 29.2 27.126.3 24.8 V2 20.9 21.5 21.4 21.9 25.4 29.7 33.0 33.0 33.6 V3 14.1 22.023.3 23.5 23.1 22.6 21.0 21.6 21.8 22.6 22.0 22.0 V4 14.0 24.1 24.2 24.126.8 22.0 23.4 24.3 24.2 24.7 25.0 23.6 M.C. 21 ± 1 27 ± 1 30 ± 1 30 ± 330 ± 2 43 45 ± 3 40 48 M.V. 17 ± 2 22 ± 1 22 ± 1 24 ± 1 26 ± 1 26 ± 2 27± 3 27 ± 3 27 ± 3 25 ± 1 24 ± 1 23 ± 1

Table 6 presents the study's results pertaining to heart rates (HR,beats per minute). During each week of the study, except for week one,it can be seen that the heart rates of the control animals (C1-C4 andS1) were higher than those of the implanted animals. Thus the inventivedevice may reduce heart rate in subjects suffering from heart failure.Put another way, the inventive device provides may enhance theefficiency of the pulmonary system and therefore reduce the frequencywith which the heart must beat to satisfy the body's oxygen demands.

TABLE 6 Heart Rate (HR, beats per minute) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk. 4Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 131 147 127 127 117 123127 143 C2 146 192 165 138 156 149 C4 135 S1 143 131 124 123 125 125 130133 131 V1 121 149 151 110 132 137 94 106 91 V2 142 132 120 140 137 144126 135 V3 151 107 74 82 111 98 95 107 112 105 96 V4 187 159 118 130 139101 72 112 122 102 M.C. 139 ± 3 157 ± 18 139 ± 13 129 ± 5 136 ± 20 133 ±8  126 ± 1  136 ± 6 133 131 M.V. 150 ± 4 137 ± 1  116 ± 16 115 ± 3 130 ±6  120 ± 12 97 ± 11 115 ± 7 108 ± 9 105 99 ± 2

Table 7 presents the study's results relating to oxygen saturation inthe vena cava (VC_SO₂, %). The control animals and the implanted animalshad similar VC_SO₂ levels throughout the course of the study, althoughfor both groups the levels were lower than at baseline. It is expectedthat oxygen saturation in the vena cava is relatively low, because thevessel carries deoxygenated blood from the body to the heart.

TABLE 7 Oxygen Saturation in Vena Cava (VC_SO₂, %) Day 0 Wk. 1 Wk. 2 Wk.3 Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 90 85 84 85 8083 80 80 79 C2 80 81 75 77 75 78 C4 82 77 62 S1 V1 94 80 80 81 79 80 6880 80 80 79 80 V2 98 78 78 70 81 78 73 79 79 V3 75 74 75 74 71 75 74 7967 74 78 V4 73 73 72 67 76 71 76 79 73 74 75 M.C. 90 82 ± 1 81 ± 2 74 ±6 79 ± 1 79 ± 4 79 ± 1 80 79 M.V. 96 ± 1 76 ± 2 76 ± 2 75 ± 2 75 ± 3 76± 2 72 ± 1 77 ± 2 79 ± 0 73 ± 4 76 ± 2 78 ± 1

Table 8 presents the study's results relating to oxygen saturation inthe pulmonary artery (PA_SO₂, %). The PA_SO₂ values for the implantedanimals are somewhat higher than those for the control animals (e.g.,between about 5-10% higher), indicating that device 100 was patent andtransferring blood from the left atrium to the right atrium. It isexpected that oxygen saturation in the pulmonary artery is relativelylow, because the vessel carries deoxygenated blood from the heart to thelungs.

TABLE 8 Oxygen Saturation in Pulmonary Artery (PA_SO₂, %) Day 0 Wk. 1Wk. 2 Wk. 3 Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 84 8176 78 71 76 75 73 C2 64 77 67 70 69 70 C4 78 76 57 S1 V1 91 81 83 82 8185 82 83 84 83 80 80 V2 92 81 80 84 87 87 80 82 84 V3 77 79 84 79 76 8078 85 71 77 81 V4 76 80 84 75 78 76 83 83 78 77 77 M.C. 84 74 ± 5 76 ± 067 ± 5 71 ± 0 69 73 ± 2 75 73 M.V. 92 ± 0 79 ± 1 81 ± 1 84 ± 1 81 ± 3 82± 3 80 ± 1 81 ± 1 84 ± 0 77 ± 3 78 ± 1 79 ± 1

Table 9 presents the oxygen saturation in the left atrium (LA_SO₂, %).The LA_SO₂ values for the implanted animals are similar to those for thecontrol animals. Animals with LA_SO₂ values of less than 94% areconsidered to have low cardiac output.

TABLE 9 Oxygen Saturation in Left Atrium (LA_SO₂, %) Day 0 Wk. 1 Wk. 2Wk. 3 Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 100 96 97 9493 95 92 96 93 C2 96 97 98 99 96 95 C4 95 95 98 S1 V1 100 93 96 97 94 9697 97 97 97 96 96 V2 100 97 97 96 92 96 87 95 97 V3 96 93 97 96 93 97 9696 94 96 96 V4 95 96 96 97 97 97 99 98 97 98 98 M.C. 100 96 ± 0 96 ± 197 ± 1 96 ± 2 96 ± 1 94 ± 1 96 93 M.V. 100 ± 0 95 ± 1 96 ± 1 97 ± 0 95 ±1 96 ± 1 95 ± 3 97 ± 1 97 ± 0 96 ± 1 97 ± 1 97 ± 1

Table 10 presents the study's results pertaining to the left ventricleinternal diameter in diastole (LVIDd, cm), which also may be referred toin the art as left ventricular end-diastolic dimension (LVEDD or LVDD).It may be seen that the LVIDd for the control (C1-C4 and S1) andimplanted (V1-V4) animals were relatively similar, and does notsignificantly vary during weeks 1-12 of the study. This may beattributed to the relatively low pressures during implantation. It maybe expected that when the device 100 is implanted in a subject with highLAP, the LVIDd will decrease after implantation as a result of thesignificant reduction in LAP.

TABLE 10 Left Ventricle Internal Diameter in Diastole (LVIDd, cm) Day 0Wk. 1 Wk. 2 Wk. 3 Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C14.6 5.4 5.0 5.1 5.4 5.3 4.8 4.8 4.8 C2 4.0 4.1 4.4 4.4 4.0 4.0 3.8 C44.2 5.7 5.7 5.5 S1 4.3 4.7 4.9 5.0 4.7 5.0 5.0 5.0 4.4 5.0 V1 3.8 4.14.2 4.3 3.8 4.0 4.1 4.5 4.3 4.4 4.3 4.0 V2 5.3 4.5 4.5 5.4 5.0 4.9 5.04.9 5.0 V3 5.4 6.3 6.2 5.9 6.0 5.6 5.5 6.0 6.2 6.3 5.9 5.6 V4 4.4 4.94.7 4.3 4.0 3.9 4.1 4.1 4.1 4.2 4.4 4.1 M.C. 4.3 ± .1 5.0 ± .4 5.0 ± .35.0 ± .2 4.7 ± .4 4.7 ± .7 4.5 ± .4 4.9 ± .1 4.9 ± .1 4.4 5.0 M.V. 4.7 ±.4 5.0 ± .5 4.9 ± .4 5.0 ± .4 4.7 ± .5 4.6 ± .4 4.7 ± .3 4.9 ± .4 4.9 ±.5 5.0 ± .7 4.9 ± .5 4.6 ± .5

Table 11 presents the study's results pertaining to the left ventricleinternal diameter in systole (LVIDs, cm), which also may be referred toin the art as left ventricular end-systolic dimension (LVESD or LVSD).While the LVIDd discussed above with respect to Table 10 was similar forboth groups of animals, it may be seen here that for the controlanimals, the LVIDs increased from baseline in week one (e.g., from anaverage 3.5±0.2 at baseline to 4.2±0.3 at week one), and then increasedfurther and/or remained elevate. By comparison, the LVIDs for theimplanted animals increased slightly from baseline in week one (e.g.,from an average 4.0±0.2 at baseline to 4.2±0.4 at week one), but thendecreased relatively steadily over the course of the study (e.g., to3.5±0.4 at week twelve). This decrease reflects the remodeling of theleft ventricle over time that results from offloading blood flow fromthe left atrium back to the right atrium through the inventive device.

TABLE 11 Left Ventricle Internal Diameter in Systole (LVIDs, cm) Day 0Wk. 1 Wk. 2 Wk. 3 Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C13.8 4.7 4.4 4.5 4.9 4.9 4.4 4.4 4.4 C2 3.0 3.3 3.8 3.8 3.5 3.7 3.6 C43.5 4.8 5.0 5.1 S1 3.6 4.1 4.3 4.4 4.2 4.5 4.6 4.6 4.7 4.7 V1 3.6 3.53.5 3.6 3.2 3.3 3.4 3.7 3.6 3.6 3.5 3.2 V2 4.7 3.8 3.7 3.8 4.0 3.9 3.93.9 4.0 V3 4.6 5.3 5.2 4.9 4.9 4.6 4.5 4.9 5.0 5.0 4.7 4.4 V4 3.4 4.03.7 3.3 3.1 2.9 3.1 3.1 3.0 3.1 3.2 2.9 M.C. 3.5 ± .2 4.2 ± .3 4.3 ± .34.5 ± .3 4.2 ± .4 4.3 ± .6 4.2 ± .3 4.5 ± .1 4.5 ± .1 4.7 4.7 M.V. 4.0 ±.3 4.2 ± .4 4.0 ± .4 3.9 ± .4 3.8 ± .4 3.7 ± .4 3.7 ± .3 3.9 ± .4 3.9 ±.4 3.9 ± .6 3.8 ± .5 3.5 ± .4

Table 12 elaborates on the results of Table 11, and presents the changesin the left ventricle internal diameter in systole (ΔLVIDs, %). As canbe seen in Table 12, the control animals experienced an average increasein LVIDs of about 20-29% over the course of the study, while theimplanted animals experienced an average decrease in LVIDs of about0-9%. Thus, the inventive device may inhibit increases in the internaldiameter of the left ventricle in subjects suffering from heart disease,and indeed may reduce the internal diameter of the left ventricle insubjects suffering from heart disease, in some embodiments by up to 10%.

TABLE 12 Change in Left Ventricle Internal Diameter in Systole (ΔLVIDs,%) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk. 4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11Wk. 12 C1 +23 +15 +18 +28 +28 +16 +16 +16 C2 +11 +25 +27 +17 +23 +20 C4+37 +43 +46 S1 +13 +17 +22 +17 +24 +26 +27 +29 +28 V1 −1 −2 +1 −11 −8 −6+4 +1 +2 −2 −10 V2 −18 −21 −19 −14 −17 −17 −17 −14 V3 +17 +13 +8 +7 +1−2 +7 +10 +10 +2 −4 V4 +19 +9 −2 −9 −12 −7 −8 −9 −8 −6 −14 M.C. +21 ± 6+25 ± 6 +28 ± 6 +21 ± 4 25 ± 2 +20 ± 2 +21 ± 5 +22 ± 6 +29 +28 M.V.  +4± 9  +0 ± 8  −3 ± 6  −7 ± 5 −9 ± 4  −8 ± 3  −4 ± 6  −3 ± 5 +1 ± 5 −2 ± 2−9 ± 3

Table 13 presents the study's results pertaining to ejection fraction(EF, %). The EF of the control animals may be seen to declinesignificantly over the course of the study, while the EF of theimplanted animals increases significantly over the course of the study.For example, it may be seen that for the control animals, C1 experienceda decline in EF to about 45% of baseline; C2 to about 28% of baseline;C4 to about 47% of baseline; and S1 to about 41% of baseline. Bycomparison, for the implanted animals, V1 experienced an increase in EFto about 169% of baseline; V2 also to about 169% of baseline; V3 toabout 129% of baseline; and V4 to about 127% of baseline. The inventivedevice thus may not only inhibit decreases in EF of subjects sufferingfrom heart failure, but indeed may increase the EF of such subjectssignificantly, for example by 25-50%, or even 25-70% or more.

TABLE 13 Ejection Fraction (EF, %) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk. 4 Wk. 5Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 35.5 28.9 26.8 23.5 21.0 18.317.8 16.4 16.0 C2 45.3 40.1 29.1 28.0 23.6 20.9 12.7 C4 34.3 32.4 25.216.2 S1 33.2 27.6 26.9 25.0 22.6 20.7 18.6 16.8 14.8 13.7 V1 24.5 27.336.1 36.6 35.9 36.0 35.7 35.7 35.6 37.7 37.8 41.4 V2 26.4 33.2 37.3 37.240.5 42.0 42.9 43.0 44.6 V3 32.6 33.6 33.3 34.5 37.2 37.2 37.9 38.2 38.941.0 41.8 41.9 V4 45.3 45.7 46.0 47.5 47.9 47.8 47.9 49.7 52.7 53.2 55.557.5 M.C. 37.1 ± 2.8 32.3 ± 2.8 27.0 ± .8  23.2 ± 2.5 22.4 ± .7  19.6 ±1.3 17.0 ± 2.3 17.5 ± 1.1 16.4 ± .4  14.8 13.7 M.V. 32.2 ± 4.7 34.9 ±3.9 38.2 ± 2.7 39.0 ± 2.9 40.4 ± 2.7 40.8 ± 2.7 41.1 ± 2.7 41.6 ± 3.142.9 ± 3.7 44.0 ± 4.7 45.0 ± 5.4 46.9 ± 5.3

Table 14 elaborates on the results presented in Table 14, and presentsthe change in ejection fraction. As can be seen in Table 14, the EF ofeach of the control animals decreased significantly relative tobaseline, e.g., by up to 72% for animal C2, while the EF for each of theimplanted animals increased significantly.

As noted above with respect to Table 10, the left ventricle internaldiameter in diastole (LVIDd) did not significantly change for theimplanted animals over the course of the study. Absent such a decreasein the LVIDd, an increase in the EF may be interpreted as an increase incardiac output. The inventive device thus may not only inhibit decreasesin cardiac output of subjects suffering from heart failure, but indeedmay increase the cardiac output of such subjects significantly.

TABLE 14 Change in Ejection Fraction (EF, %) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk.4 Wk. 5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 −18 −24 −34 −41 −48−50 −54 −55 C2 −11 −36 −38 −48 −54 −72 C4 −6 −19 −53 S1 −17 −27 −25 −32−38 −44 −49 −55 −59 V1 +11 +47 +49 +46 +47 +46 +45 +45 +54 +54 +69 V2+26 +42 +41 +54 +59 +63 +63 +69 V3 +3 +2 +6 +14 +14 +16 +17 +19 +26 +28+29 V4 +1 +2 +5 +6 +6 +6 +10 +16 +18 +23 +27 M.C. −13 ± 3 −26 ± 4 −37 ±6  −40 ± 5  −51 ± 3  −53 ± 10 −49 ± 5  −52 ± 3  −55 −59 M.V. +10 ± 6 +23± 2 +25 ± 12 +30 ± 12 +32 ± 13 +33 ± 13 +34 ± 12 +38 ± 12 +32 ± 11 +35 ±10 +41 ± 14

Table 15 presents the study's results pertaining to fractionalshortening (FS, %). Similar to ejection fraction discussed above withrespect to Tables 13-14, the FS of each of the control animals may beseen in Table 15 to decline significantly over the course of the study.For example, animal C1 experienced a decline in FS to about 47% ofbaseline before death; animal C2 to about 24% of baseline; animal C4 toabout 46% of baseline; and animal S1 to about 39% of baseline. Incontrast, the FS of each of the implanted animals increasedsignificantly over the course of the study. For example, animal V1experienced an increase in FS to about 183% of baseline; animal V2 toabout 166% of baseline; animal V3 to about 132% of baseline; and animalV4 to about 127% of baseline. Thus, the inventive device not onlyinhibits decreases in fractional shortening for subjects suffering fromheart failure, but also may increase fractional shorteningsignificantly, e.g., by about 25-85% of baseline.

TABLE 15 Fractional Shortening (FS, %) Day 0 Wk. 1 Wk. 2 Wk. 3 Wk. 4 Wk.5 Wk. 6 Wk. 8 Wk. 9 Wk. 10 Wk. 11 Wk. 12 C1 17.0 13.7 12.5 10.9 9.7 8.48.0 7.5 8.0 C2 23.2 19.3 13.5 13.0 10.7 9.1 5.5 C4 16.2 15.5 11.8 7.4 S115.6 12.8 12.5 11.6 10.3 9.4 8.4 7.6 6.6 6.1 V1 10.9 12.6 17.1 17.5 16.916.9 16.9 17.0 16.9 18.1 17.6 20.0 V2 12.4 15.8 18.1 19.0 19.9 20.7 21.221.6 20.6 V3 15.7 16.4 16.2 16.7 18.3 18.2 18.5 18.8 19.3 20.5 20.8 20.8V4 22.4 22.6 22.9 23.7 23.7 23.6 23.8 24.9 26.7 27.1 28.8 28.4 M.C. 18.0± 1.8 15.3 ± 1.4 12.6 ± 0.4 10.7 ± 1.2 10.2 ± 0.3  8.7 ± 0.4  7.7 ± 1.2 8.0 ± 0.4  7.8 ± 0.2 6.6 6.1 M.V. 15.3 ± 2.5 16.8 ± 2.1 18.6 ± 1.5 19.2± 1.6 19.7 ± 1.5 19.8 ± 1.5 20.1 ± 1.5 20.6 ± 1.7 20.9 ± 2.1 21.9 ± 2.722.4 ± 3.3 23.1 ± 2.7

As the foregoing results illustrate, devices constructed and implantedaccording to the present invention may provide for significantlyimproved mortality rates in subjects suffering from heart failure. Inparticular, the devices may significantly enhance ejection fraction,fractional shortening, and/or cardiac output in subjects who wouldotherwise have significantly diminished cardiac function as a result ofexcessive left atrial and left ventricular pressures. For example,subjects may be classified under the New York Heart Association (NYHA)classification system as having Class II (Mild) heart failure, who haveslight limitation of physical activity and are comfortable at rest, butfor whom ordinary physical activity results in fatigue, palpitation, ordyspnea; Class III (Moderate) heart failure, who have marked limitationof physical activity, may be comfortable at rest, and may experiencefatigue, palpitation, or dyspnea if they engage in less than normalactivity; or as having Class IV (Severe) heart failure, who are unableto carry out any physical activity without discomfort, exhibit symptomsof cardiac insufficiency at rest, and have increased discomfort if theyundertake any physical activity. The present devices may significantlyincrease the cardiac output of such class III or class IV subjects,particularly those with low ejection fraction, enabling them to engagein significantly more physical activity than they otherwise could. Thepresent devices further may decrease pulmonary artery pressure insubjects with left heart failure, and additionally may reduce or inhibitpulmonary congestion in patients with pulmonary congestion resultingfrom such heart failure, for example by inhibiting episodes of acutepulmonary edema. Indeed, as the above-described Example illustrates, theinventive device may reduce LAP and PAP significantly relative to whatthose pressures would otherwise be; such pressure reductions may notonly provide immediate relief from acute symptoms, but further mayfacilitate cardiac remodeling over the weeks following implant and thusprovide for enhanced cardiac function. The devices may in someembodiments include means for measuring the various parameters ofinterest, e.g., means such as discussed above with respect to the animaltrials.

It should be noted that the inventive devices also may be used withpatients having disorders other than heart failure. For example, in oneembodiment the device may be implanted in a subject suffering frommyocardial infarction, for example in the period immediately followingmyocardial infarction (e.g., within a few days of the event, or withintwo weeks of the event, or even within six months of the event). Duringsuch a period, the heart remodels to compensate for reduced myocardialfunction. For some subjects suffering from severe myocardial infarction,such remodeling may cause the function of the left ventricle tosignificantly deteriorate, which may lead to development of heartfailure. Implanting an inventive device during the period immediatelyfollowing myocardial infarction may inhibit such deterioration in theleft ventricle by reducing LAP and LVEDP during the remodeling period.For example, in the above-described Example, heart failure was inducedin the sheep by injecting microspheres that block the coronary arteryand induce myocardial infarction. Following the myocardial infarction,the sheep developed heart failure. As can be seen in the various resultsfor the implanted animals, implanting the inventive device even a weekfollowing the myocardial infarction inhibited degradation of the heartand yielded significantly improved mortality rates and cardiacfunctioning both immediately and over time as the subjects' heartsremodeled. As such, it is believed that implanting an inventive devicefor even a few weeks or months following myocardial infarction mayprovide significant benefits to the subject as their heart remodels. Thedevice optionally then may be removed.

While various illustrative embodiments of the invention are describedabove, it will be apparent to one skilled in the art that variouschanges and modifications may be made herein without departing from theinvention. It will further be appreciated that the devices describedherein may be implanted in other positions in the heart. For example,device 100 illustrated in FIGS. 1A-1D may be implanted in an orientationopposite to that shown in FIG. 2B, so as to shunt blood from the rightatrium to the left atrium, thus decreasing right atrial pressure; such afeature may be useful for treating a high right atrial pressure thatoccurs in pulmonary hypertension. Similarly, device 100 may be implantedacross the ventricular septum, in an orientation suitable to shunt bloodfrom the left ventricle to the right ventricle, or in an orientationsuitable to shunt blood from the right ventricle to the left ventricle.The appended claims are intended to cover all such changes andmodifications that fall within the true spirit and scope of theinvention.

What is claimed:
 1. A method of delivering a shunt to an atrial septumof a patient to reduce atrial pressure of the patient, the methodcomprising: selecting an atrial shunt having first and second conicalsections, each of the first and second conical sections having a flaredend region, wherein apices of the first and second conical sections arejoined together to form a neck region disposed in a mid-region of theatrial shunt, the atrial shunt defining a shunt lumen configured toshunt blood through the atrial septum to reduce atrial pressure;selecting a sheath configured for transvascular insertion via thepatient's blood vessel and a pusher configured to be disposed in a lumenof the sheath; loading the atrial shunt into the sheath, wherein thefirst and second conical sections are compressed inward relative to alongitudinal axis of the shunt to a contracted position, and a proximalend of the second conical section is disposed in contact with thepusher; percutaneously delivering the sheath through an opening formedin the atrial septum such that a distal region of the sheath is disposedwithin a left atrium; advancing the pusher distally relative to thesheath to expose the first conical section so that it transitions from acompressed state to an expanded state within the left atrium; retractingthe shunt and the sheath proximally such that the first conical sectionengages the atrial septum from within the left atrium and the neckregion engages the opening in the atrial septum, thereby enabling theshunt to self-locate in the opening; and further retracting the sheathproximally such that the first conical section flanks the atrial septumfrom within the left atrium, the neck region is lodged in the openingformed in the atrial septum, and the second conical section is exposedfrom the sheath and transitions from a compressed state to an expandedstate within the right atrium, thereby facilitating shunting of bloodthrough the shunt lumen between the left atrium and the right atrium. 2.The method of claim 1, further comprising mechanically locking thepusher such that further advancement of shunt within the lumen of thesheath is prevented.
 3. The method of claim 1, wherein percutaneouslydelivering comprises percutaneously delivering the sheath via a femoralvein access point.
 4. The method of claim 1, further comprising, priorto percutaneously delivering the sheath, percutaneously delivering atleast one of a needle and a guidewire to the atrial septum andpuncturing the atrial septum.
 5. The method of claim 4, wherein thesheath is delivered over the guidewire.
 6. The method of claim 1,further comprising assessing the efficacy of the shunt via fluoroscopicvisualization of the shunt.
 7. The method of claim 1, wherein the shuntcomprises an hourglass configuration when the first and second conicalsections are in the expanded state.
 8. The method of claim 1, whereinthe second conical section is configured to flank the atrial septum fromwithin the right atrium in the expanded state.
 9. The method of claim 1,further comprising removing the sheath from the patient.